WO2007043137A1 - Nuclear medical diagnosis device - Google Patents

Nuclear medical diagnosis device Download PDF

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Publication number
WO2007043137A1
WO2007043137A1 PCT/JP2005/018360 JP2005018360W WO2007043137A1 WO 2007043137 A1 WO2007043137 A1 WO 2007043137A1 JP 2005018360 W JP2005018360 W JP 2005018360W WO 2007043137 A1 WO2007043137 A1 WO 2007043137A1
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WO
WIPO (PCT)
Prior art keywords
timing
scintillator
incident
light
nuclear medicine
Prior art date
Application number
PCT/JP2005/018360
Other languages
French (fr)
Japanese (ja)
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WO2007043137A9 (en
Inventor
Jyunichi Ooi
Original Assignee
Shimadzu Corporation
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Application filed by Shimadzu Corporation filed Critical Shimadzu Corporation
Priority to JP2007539758A priority Critical patent/JPWO2007043137A1/en
Priority to PCT/JP2005/018360 priority patent/WO2007043137A1/en
Priority to US12/088,231 priority patent/US7709801B2/en
Priority to CN2005800502280A priority patent/CN101208616B/en
Publication of WO2007043137A1 publication Critical patent/WO2007043137A1/en
Publication of WO2007043137A9 publication Critical patent/WO2007043137A9/en

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Classifications

    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
    • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
    • G01T1/16—Measuring radiation intensity
    • G01T1/161—Applications in the field of nuclear medicine, e.g. in vivo counting
    • G01T1/164—Scintigraphy
    • G01T1/1641—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras
    • G01T1/1644—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras using an array of optically separate scintillation elements permitting direct location of scintillations

Definitions

  • a radiopharmaceutical is administered to a subject, and a pair of ⁇ -rays emitted from a positron radioisotope (radioisotope, RI) accumulated in a region of interest of the subject is simultaneously measured, and the subject is interested.
  • the present invention relates to a nuclear medicine diagnostic apparatus (ECT apparatus) for obtaining a tomographic image of a region, and in particular, to a technique for simultaneously counting ⁇ rays.
  • ECT apparatus nuclear medicine diagnostic apparatus
  • a nuclear medicine diagnosis apparatus that is, an ECT (Emission Computed Tomography) apparatus
  • a PET (Positron Emission Tomography) apparatus will be described as an example.
  • the PET device detects the force of the region of interest of the subject by detecting the two gamma rays emitted in a direction of approximately 180 ° from each other with the opposing gutter detectors, and when these gamma rays are simultaneously detected (simultaneously counted). It is configured to reconstruct a tomographic image of the specimen.
  • the ⁇ -ray detector used to simultaneously count ⁇ -rays with a PET device includes a scintillator that emits light when ⁇ -rays emitted from the subject force enter, and converts the light emitted from the scintillator into an electrical signal. Some are composed of photomultiplier tubes.
  • the scintillator is divided (optically coupled) into scintillators with different light emission attenuation times in the direction of ⁇ -ray incidence, for example, the scintillator has a short ⁇ -ray attenuation time on the ⁇ -ray incidence side,
  • the ⁇ -ray decay time is divided into a scintillator array on the photomultiplier tube side, and the position of the emitted ⁇ -ray is detected accurately even when the ⁇ -ray is obliquely incident on the scintillator of the ⁇ -ray detector. Improvements are made to obtain more accurate tomographic images (see, for example, Patent Documents 1 and 2).
  • Patent Document 1 JP-A-6-337289 (Page 2-3, Fig. 1)
  • Patent Document 2 JP 2000-56023 A (Page 2-3, Fig. 1)
  • the conventional nuclear medicine diagnostic apparatus has the following problems.
  • scintillators with different emission pulse decay times often have different emission pulse rise times. Therefore, when simultaneous counting is performed by a ⁇ -ray detector using a scintillator array with different rise times of the light emission pulses, a time difference occurs in detection between the opposing ⁇ -ray detectors. In other words, due to this time difference, ⁇ rays that are emitted simultaneously are not recognized as being simultaneously emitted in the coincidence process based on the detection by the ⁇ detector, and the detection sensitivity decreases. There is a problem of making it.
  • the time range (timing window) for effective coincidence counting in the coincidence counting process is widened, and there is a time difference in detection between opposing ⁇ -ray detectors. Even if it is determined to be simultaneous, there is a problem that the influence of accidental coincidence counting and scattering coincidence counting increases, leading to degradation of the reconstructed image.
  • the present invention has been made in view of such circumstances, and even when a scintillator having a different emission pulse decay time is used for the ⁇ -ray detector, the sensitivity is high, and An object of the present invention is to provide a nuclear medicine diagnostic apparatus capable of obtaining an accurate tomographic image without causing deterioration.
  • the present invention has the following configuration.
  • a plurality of scintillators are two-dimensionally closely arranged, and a plurality of scintillator arrays having different emission pulse decay times in the y-ray incident depth direction are optically coupled.
  • the incident Timing calculation hand It is determined whether or not to correct the incident timing calculated in the stage, and incident timing correction means for correcting the incident timing based on the determination result. It is.
  • the operation of the invention of claim 1 is as follows.
  • the ⁇ rays from which the subject force was also released are a two-dimensional arrangement of multiple scintillators, and optically combined multiple scintillator arrays with different emission pulse decay times in the direction of the ⁇ -ray incident depth. Incident into the scintillation tab lock.
  • the ⁇ -rays incident on the scintillator block emit light by each scintillator of the scintillator array having a different emission pulse decay time.
  • the light emission light emitted by each scintillator is converted into an electric signal by the light receiving element.
  • the incident timing calculation means calculates the timing at which the light signal is also incident on the scintillator array for the output electrical signal.
  • the scintillator array identifying means identifies which scintillator array among a plurality of electrical signals output from the light receiving elements is incident.
  • the incident timing correction means determines whether or not to correct the incident timing calculated by the incident timing calculation means in accordance with the scintillator array identified by the scintillator array identification means. Based on the above, the incident timing is corrected. Therefore, the incident timing correction means determines whether or not to correct the incident timing calculated by the incident timing calculation means according to the scintillator array identified by the scintillator array identification means, and based on the determination result.
  • the incident timing is corrected, even when simultaneous counting is performed using scintillator arrays with different emission pulse decay times, the detection time difference between the scintillator arrays caused by the different decay times of the emission pulses can be reduced. It can be solved by correction. Therefore, it is possible to increase the detection sensitivity and obtain an accurate tomographic image without causing deterioration of the reconstructed image.
  • the nuclear medicine diagnosis apparatus of the invention of claim 2 includes an AZD modification that converts an analog signal that is an electric signal output from the light receiving element into a digital signal
  • the scintillation array identifying means includes an AZD modification.
  • the addition means for sequentially adding the digital signals converted in step 1 and in the addition means, in the middle of adding the digital signals from the start of light emission emitted by the scintillator block to the middle of the light emission.
  • Addition value Identification for calculating the identification value indicating the value obtained by dividing the intermediate addition value by the total addition value from the total addition value obtained by adding the digital signals up to the end of light emission of the emission pulse emitted by the scintillator block It is determined whether the identification value calculated by the identification value calculation means is a large threshold value or a small threshold value with respect to an intermediate value between the identification values of the scintillator array calculated by the value calculation means and the identification value calculation means. It is characterized by having a discrimination means.
  • the AZD conversion converts an analog signal that is an electrical signal output from the light receiving element into a digital signal.
  • the addition means of the scintillator array identification means sequentially adds the digital signals converted by the AZD converter.
  • the identification value calculation means adds the digital signal up to the midpoint of the light emission start power of the light emission pulse emitted by the scintillator block obtained by addition by the addition means until the end of light emission. Identification that shows the value obtained by dividing the intermediate addition value by the half-calculated value from the addition value and the total addition value obtained by adding the digital signals until the end of light emission of the light emission pulse emitted by the scintillator block Calculate the value.
  • the intermediate value calculation means obtains an intermediate value between the identification values of each scintillator array calculated by the identification value calculation means, and the determination means determines the intermediate value calculated by the intermediate value calculation means. It is determined whether the discriminant value calculated by the discriminant value calculation means is a large value vj or a small value. Therefore
  • the scintillator array identifying means can identify which scintillator array is the light emission pulse emitted by the scintillator.
  • the integration operation that has been conventionally performed by the integrator can be replaced with an addition operation in which the addition means sequentially adds, the number of parts can be reduced and the cost can be reduced.
  • the nuclear medicine diagnosis apparatus of the invention of claim 3 performs simultaneous counting using the incident timing corrected by the incident timing correcting means and the incident timing determined not to correct the incident timing by the incident timing correcting means.
  • a timing window storage means for storing a timing window indicating a predetermined range of the same clock number by the simultaneous counting means as a timing window corresponding to each combination of a plurality of scintillator arrays It is characterized by having these.
  • the coincidence counting unit calculates the incident timing corrected by the incident timing correcting unit and the incident timing determined not to correct the incident timing by the incident timing correcting unit. To perform coincidence.
  • the coincidence counting is performed using a timing window indicating a predetermined range in which the coincidence counting is determined to be simultaneous, and here, it corresponds to a combination of a plurality of scintillator arrays stored in the timing window storage means.
  • the timing window is used for simultaneous counting. Therefore, by using different timing windows depending on the combination of each of the multiple scintillator arrays, it is possible to perform highly accurate coincidence, reduce the effects of accidental coincidence and scattered clocks, reduce noise, and achieve high image quality. Can be obtained.
  • the nuclear medicine diagnosis apparatus of the invention of claim 4 performs simultaneous counting using the incident timing corrected by the incident timing correcting means and the incident timing determined not to correct the incident timing by the incident timing correcting means.
  • a timing window storage means for storing a timing window indicating a predetermined range of the same clock number by the simultaneous counting means as a timing window corresponding to each combination of a plurality of scintillators. It is characterized by that.
  • the coincidence counting unit calculates the incident timing corrected by the incident timing correcting unit and the incident timing determined not to correct the incident timing by the incident timing correcting unit. To perform coincidence.
  • the coincidence counting is performed using a timing window indicating a predetermined range in which the coincidence counting is determined to be simultaneous.
  • a timing window corresponding to each combination of a plurality of scintillators stored in the timing window storage means. Are simultaneously counted. Therefore, by using different timing windows depending on the combination of multiple scintillators, it is possible to perform highly accurate coincidence counting, reduce the effects of accidental coincidence counting and scattering coincidence counting, reduce noise and achieve high image quality. An image can be obtained.
  • the nuclear medicine diagnosis apparatus of the invention of claim 5 is characterized by comprising a light guide for optically coupling the scintillator block and the light receiving element.
  • the scintillator block and the light receiving element And a light guide that optically couples the two. Therefore, the light guide can appropriately guide the light from the scintillator block to the light receiving element.
  • the plurality of scintillator arrays are made of Gd SiO (GSO) having a Ce concentration of 0.5 mol, Gd SiO (GSO), Lu having a Ce concentration of 1.5 mol. SiO (LS
  • the plurality of scintillator arrays are made of Gd SiO (GSO) having a Ce concentration of 0.5 mol, Gd SiO (GSO), Lu having a Ce concentration of 1.5 mol. SiO
  • the nuclear medicine diagnosis apparatus of the invention of claim 7 is characterized in that the light receiving element is a photomultiplier tube.
  • the light receiving element is a photomultiplier tube
  • the light from the scintillator block can be appropriately converted into an electric signal.
  • the nuclear medicine diagnosis apparatus of the invention of claim 8 is characterized in that the light receiving element is a photodiode.
  • the light receiving element is a photodiode
  • the light from the scintillator block can be appropriately converted into an electric signal.
  • the nuclear medicine diagnosis apparatus of the invention of claim 9 is characterized in that the light receiving element is an avalanche photodiode.
  • the light receiving element is an avalanche photodiode, it is possible to appropriately convert the light from the scintillator block into an electric signal.
  • the incident timing correction unit corrects the incident timing calculated by the incident timing calculation unit according to the scintillator array identified by the scintillator array identification unit. Since the incident timing is corrected based on the result of the determination, the decay time of the light emission pulse can be obtained even when simultaneous counting is performed using a scintillator array having a different decay time of the light emission pulse. The time difference of detection between the scintillator arrays caused by the difference between the two can be eliminated by correction. Therefore, it is possible to increase the detection sensitivity and obtain an accurate tomographic image without causing deterioration of the reconstructed image.
  • FIG. 1 is a block diagram showing an overall configuration of a PET apparatus.
  • FIG. 2 is a block diagram showing a configuration of an FPGA.
  • FIG. 3 is a perspective view showing a configuration of a y-ray detector.
  • FIG. 4 is a graph showing light emission pulses of each scintillator array output from the amplifier circuit.
  • FIG. 5 (a) and (b) are diagrams showing the timing of ⁇ rays incident on each scintillator array.
  • FIG. 6 is a graph showing the addition power until the end of light emission as well as the light emission start force of the light emission pulse.
  • FIG. 7 is a graph for explaining a timing window.
  • FIG. 8 is a graph showing the timing spectrum when V is not corrected for the difference in detection time due to different decay times of the scintillator array.
  • FIG. 9 is a graph showing a timing spectrum when a difference in detection time due to different decay times of the scintillator array is corrected.
  • FIG. 10 is a flowchart showing signal processing after the amplification circuit power is also output.
  • Incident timing calculator (incident timing calculator)
  • Scintillator array identification unit (scintillator array identification means)
  • Identification value calculation unit (identification value calculation means)
  • Incident timing correction unit (incident timing correction means)
  • Timing window storage (timing window storage means)
  • FIG. 1 is a block diagram showing the overall configuration of the PET apparatus.
  • FIG. 2 is a block diagram showing the configuration of FPGA7.
  • a PET apparatus will be described as an example of a nuclear medicine diagnosis apparatus.
  • the PET device receives a radiopharmaceutical to subject M and enters ⁇ -rays emitted from positron radioisotopes (radioisotopes, RI) accumulated in the region of interest of subject M.
  • a ⁇ -ray detector 1 that generates light, converts the light into an electrical signal, and outputs the signal is provided.
  • the X-ray detector 1 is arranged around the body axis of the subject X, for example, in a ring shape with a diameter of about 700 mm without any gap (in FIG. 1, only two ⁇ -ray detectors 1 are shown). Yes.
  • the two gamma rays emitted from the subject's region of interest in the direction of approximately 180 ° are detected by the opposing shoreline detector 1, converted into an electrical signal, and output.
  • amplifier circuits 2a and 2b that amplify the electrical signal output from the ⁇ -ray detector 1, and an AZD converter 3a that converts the analog signal amplified by the amplifier circuits 2a and 2b into a digital signal.
  • the two ⁇ -ray detectors 1 include a coincidence processing unit 6 that performs processing for detecting (simultaneous counting) that ⁇ rays are incident simultaneously. I have.
  • the position calculation processing unit 5 and the coincidence counting processing unit 6 are provided in a programmable LSI (Large Scale Integrated Circuit) called an FPGA (Field Programmable Gate Array) 7. ing.
  • the FPGA 7 has functions such as CPU 8, ROM 9, and RAM 10, and the position calculation processing unit 5 and the coincidence processing unit 6 are functions of the CPU 8 of the FPGA 7.
  • the reconstruction unit 11 reconstructs a tomographic image of the subject. Yes.
  • the apparatus of this embodiment includes a controller 12, a monitor 13, an input unit 14, and the like.
  • a controller 12 controls the apparatus of this embodiment.
  • monitor 13 controls the apparatus of this embodiment.
  • input unit 14 controls the apparatus of this embodiment.
  • FIG. 3 is a perspective view showing the configuration of the ⁇ -ray detector 1.
  • the ⁇ -ray detector 1 is a DOI (Depth Of lnteraction) detector in which the scintillator 19 is divided and arranged in the ⁇ -ray incident depth direction, that is, the scintillator is arranged three-dimensionally. is there.
  • this DOI detector is composed of a scintillator block 15, a light guide 16, and a photomultiplier tube (PMT) 17.
  • PMT photomultiplier tube
  • the scintillator block 15 is an optically coupled two scintillator array 18a and scintillator array 18b having different emission pulse decay times in the ⁇ -ray incident depth direction (Z direction).
  • 18a is a plurality of scintillators 19a
  • scintillator array 18b is a two-dimensional arrangement of scintillators 19b.
  • the scintillator block 15 includes a scintillator array 18a using a scintillator 19a (for example, Lu Y SiO (LYSO)) with a short decay time of the emission pulse on the ⁇ -ray incident side (front stage).
  • LYSO Lu Y SiO
  • a scintillator array 18b that uses a scintillator 19b (for example, Gd SiO (GSO) with a Ce concentration of 0.5 mol) on the light guide 16 side (rear stage) with a long decay time of the emission light.
  • a scintillator 19b for example, Gd SiO (GSO) with a Ce concentration of 0.5 mol
  • the two scintillator arrays 18a and 18b are each composed of 8 ⁇ 8 chip-shaped scintillators 19a and 19b (X direction and Y direction), and adjacent scintillators 19a and 18b in the scintillator arrays 18a and 18b.
  • a light reflecting material for proportionally distributing the light generated by the incidence of ⁇ rays in the X and ⁇ directions are inserted or filled depending on the location.
  • the light guide 16 guides light generated by the scintillators 19 a and 19 b of the scintillator block 15 to the photomultiplier tube 17, and is interposed between the scintillator block 15 and the photomultiplier tube 17. Inserted and each optically bonded with an optical adhesive.
  • the photomultiplier tube 17 has, for example, four photoelectric conversion films (channels) built in, and the light generated by the scintillators 19a and 19b is incident on the four-surface PMT photoelectric conversion films and is electronically amplified. After that, it is finally converted into an electrical signal (analog signal) and output. Therefore, the output from the photomultiplier tube 17 becomes the output of the ⁇ -ray detector 1.
  • the photomultiplier tube 17 described above corresponds to a light receiving element.
  • FIG. 4 is a graph showing light emission pulses of the scintillator arrays 18a and 18b output from the amplifier circuit 2a or the amplifier circuit 2b.
  • FIGS. 5 (a) and 5 (b) are diagrams showing the timing of ⁇ rays incident on the scintillator arrays 18a and 18b.
  • the curve ( ⁇ ) shown in Figs. 4 and 5 (a) shows the light incident on the scintillator array 18a with a short decay time of the emission pulse, and the curve (B) shows the decay time of the emission pulse. Shown incident on long scintillator array 18b.
  • FIG. 4 is a graph showing light emission pulses of the scintillator arrays 18a and 18b output from the amplifier circuit 2a or the amplifier circuit 2b.
  • FIGS. 5 (a) and 5 (b) are diagrams showing the timing of ⁇ rays incident on the scintillator arrays 18a and 18b.
  • the electrical signals output from the ⁇ -ray detector 1 are input to the incident timing calculation units 4a and 4b via the amplifier circuits 2a and 2b, and ⁇ -rays are generated based on the electrical signals.
  • the incident timing calculation units 4 a and 4 b include a wave height and rise time compensation circuit called ARC (Amplitude and Rise-time Compensation) 20 and a timing generation circuit 21.
  • ARC Amplitude and Rise-time Compensation
  • the ARC 20 based on the ⁇ -rays incident on the scintillator arrays 18a and 18b of the ⁇ -ray detector 1, for example, attenuation of the light emission pulses output from the amplifier circuits 2a and 2b as shown in FIG. Analog signals with different times are input. Further, the ARC 20 performs a waveform shaping process for calculating the incident timing of the ⁇ rays incident on the scintillator arrays 18a and 18b for each of these analog signals. Specifically, in this waveform shaping process, the signal obtained from the amplification circuits 2a and 2b is delayed and the voltage value of the signal obtained from the amplification circuits 2a and 2b is inverted and lowered.
  • t and t indicate the incident timing of the ⁇ -rays incident on the scintillator arrays 18a and 18b.
  • the timing generation circuit 21 converts the signal indicating the incident timing calculated by the ARC 20 as shown in FIG. 5 (b) into a digital signal, and stores it in the incident timing storage unit 22 which is one function of the RAM 10 of the FPGA 7. The configuration is temporarily stored.
  • the incident timing calculation units 4a and 4b described above correspond to incident timing calculation means.
  • the position calculation processing unit 5 will be described with reference to FIG. First, as shown in FIG. 1, the electrical signal output from the ⁇ -ray detector 1 is input via the amplifier circuits 2a and 2b, and the analog signal input from the amplifier circuits 2a and 2b is always converted to AZD. AZD conversion to be converted The digital signal power converted by the 3a, 3b is temporarily stored in the AZD conversion signal storage unit 23 which is a function of the RAM 7 of the FPGA 7. Based on the digital signal stored in the AZD conversion signal storage unit 23, the position calculation processing units 5a and 5b detect the positions of the scintillators 19a and 19b of the ⁇ -ray detector 1 on which the ⁇ -rays emitted from the subject M are incident.
  • the position calculation processing units 5a and 5b are configured to detect ⁇ detected by the ⁇ -ray detector 1.
  • a scintillator array identifying unit 24 for identifying which of the two line forces is detected based on the incident on the scintillator arrays 18a and 18b is provided. In other words, when this scintillator array 18a or scintillator array 18b is identified, the position is calculated in the Z direction indicating whether it is the scintillator 19a or scintillator 19b of the winding detector 1. Will be.
  • FIG. 6 is a graph showing the added value from the start of light emission to the end T of light emission. Figure 6 shows
  • the curve in (A) shows that the emission pulse has a short decay time and is incident on the scintillator 19a (scintillator array 18a).
  • the curve in (B) shows the scintillator 19b (scintillator 19b in which the emission pulse has a long decay time. Shown incident on array 18b).
  • An adder 25 that sequentially adds the digital signals converted by the A / D converters 3a and 3b, and in this adder 25, the light emission starting power of the light emitting light emitted by the scintillators 19a and 19b is also until the light emission end T.
  • T2 T1 is divided by the total added value A).
  • Identification value calculation unit 26 for the intermediate value K between the identification values of each scintillator array calculated by the scintillator array 26 based on the emission pulses emitted by the scintillators 19a and 19b of the modulator array
  • a value vj having a large value and a determination unit 28 for determining whether the value is a small value are provided. Therefore, based on the discrimination result in the discriminator 28, it is determined which of the two ⁇ -ray forces detected by the ⁇ -ray detector 1 is detected based on the incident light on the scintillator arrays 18a and 18b. It can be identified. Calculate A / A
  • the intermediate value K is the value A at the time Fs X m (Fs X m: Fs is the sampling interval of AZD conversion, m is the number of additions) of the two patterns of waveforms in the addition process in the adder 25. Set both wave heights
  • the scintillator array identifying unit 24 described above corresponds to a scintillator array identifying unit.
  • the adding unit 25 described above corresponds to adding means.
  • the identification value calculation unit 26 described above corresponds to an identification value calculation unit.
  • the determination unit 28 described above corresponds to a determination unit, and reads the intermediate value stored in the intermediate value data table 27 during the determination process.
  • the position calculation processing units 5a and 5b correct the incident timings calculated by the incident timing calculation units 4a and 4b according to the scintillator arrays 18a and 18b identified by the scintillator array identification unit 24.
  • An incident timing correction unit 29 is provided for determining whether or not, and correcting the incident timing based on the determination result. Specifically, when the scintillator of the scintillator array identified by the scintillation array identifying unit 24 is identified as the scintillator 19b having a long decay time, the incident timing calculated by the incident timing calculating units 4a and 4b Execute t processing where t is t — ⁇ t (incident timing correction value)
  • the corrected incident timing storage unit 30 which is a function of the RAM 10. Conversely, when the scintillator of the scintillator array identified by the scintillator array identifying unit 24 is identified as the scintillator 19a having a short decay time, the incident timing t is not corrected,
  • the corrected incident timing storage unit 30 includes a scintillator array 18a and a scintillator in relation to the incident timing t and the incident timing t.
  • the structure is such that the incident timing t and the incident timing t, in which the difference in detection time due to the difference in attenuation time from the array 18b is corrected, are temporarily stored.
  • the irradiation timing correction unit 29 corresponds to incident timing correction means.
  • This ⁇ t (incidence timing correction value) is obtained in advance by experiment as data for performing correction, and the time difference between rise times different between the scintillator array 18a and the scintillator array 18b is obtained.
  • As an incident timing correction value it is stored in a correction data table 31 which is one function of ROM9 of FPGA7.
  • the incident timing correction unit 29 reads out the incident timing correction value stored in the correction data table 31 during the correction process.
  • the coincidence processing unit 6 includes the incident timing stored in the corrected incident timing storage unit 30. t and t are read at regular intervals (for example, 128 ns), and two incident timings from ⁇ -ray detector 1 are read.
  • T, t in this case, four combinations are counted simultaneously, and these four combinations are counted.
  • a valid coincidence count is set. Otherwise, the count is invalid.
  • FIG. 7 is a graph for explaining the timing window Tw. If the vertical axis A represents the number of events (y-line detection for simultaneous counting) and the horizontal axis T represents the time difference for the detection of ⁇ -rays, the timing spectrum shown in Fig. 7 is obtained. The timing spectrum shows that when the horizontal axis 0 is 0 (there is no time difference in the detection of ⁇ rays), the number of events increases. The longer the time difference, the smaller the number of events. In other words, the horizontal axis ⁇ is 0.
  • the graph is close to a Gaussian distribution with the peak number of events on the vertical axis ⁇ ⁇ .
  • an intermediate value AZ2 of the vertical axis A in the timing spectrum shown in FIG. 7 is a half width, and a time range obtained by doubling the half width is a timing window Tw.
  • FIG. 8 is a graph showing the timing spectrum when the difference in detection time due to the different decay times of the scintillator array is not corrected.
  • Figure 9 is a graph showing the timing spectrum when the difference in detection time due to different decay times of the scintillator array is corrected.
  • the coincidence processing unit 6 Since the difference in detection time due to the difference in the decay time between the scintillator array 18a and the scintillator array 18b is not corrected, the coincidence processing unit 6 The four simultaneous counts performed are, for example, MDlt and MDlt when one ⁇ -ray detector 1 of two ⁇ -ray detectors 1 is MD1, and the other ⁇ -ray detector 1 is MD2.
  • the difference in detection time due to the difference in attenuation time between the scintillator array 18a and the scintillator array 18b is corrected by the incident timing correction unit 29, which is shown in FIG.
  • the four types of coincidence performed by the coincidence processing unit 6 are within the timing window Tw, and the effective coincidence count is not counted down, and the sensitivity does not decrease.
  • the timing window Tw is stored in the timing window storage unit 32 which is a function of the ROM 9 of the FPGA 7.
  • the above-described coincidence processing unit 6 corresponds to coincidence means.
  • the timing window storage unit 32 described above corresponds to a time window storage unit.
  • FIG. 10 is a flow chart from the incident timing generation to the incident timing correction processing.
  • the y-line represents the scintillator array 18a of the scintillator array 18a with a short decay time of the light emission pulse constituting the scintillator block 15, and the scintillator 19b of the scintillator array 18b with a long decay time of the light emission pulse.
  • the photomultiplier tube (PMT) 17 is based on the incident position (X direction and Y direction of the scintillators 19a and 19b). It is distributed to the photoelectric conversion film. Further, in the photomultiplier tube 17, the light is converted into an electric signal (analog signal) and output to the amplification circuits 2a and 2b. In the amplification circuits 2a and 2b, the analog signal is voltage amplified and output to the incident timing calculation units 4a and 4b and the AZD variables 3a and 3b. Further, the analog signal input to the AZD transformations 3 & , 3b is AZD converted into a digital signal and temporarily stored in the AZD conversion signal storage unit 23.
  • the amplification circuits 2a and 2b Ana mouth from The incident timings t and t are calculated by the incident timing calculators 4a and 4b,
  • the flow until the incident timing t, t is processed by the incident timing correction unit 29 is as follows.
  • FIG. 10 is a flowchart showing signal processing after the electrical signal output from the photomultiplier tube 17 is output from the amplifier circuits 2a and 2b.
  • Step S1 The ARC 20 of the incident timing calculation units 4a and 4b calculates the incident timings t and t based on the analog signals from the amplifier circuits 2a and 2b, and further generates the timing.
  • the live circuit 21 converts these incident timings t and t into digital signals and records the incident timings.
  • step SI Until t is generated, the next operation is not performed (the operation of step SI is repeated).
  • Step S2 The scintillator array identification unit 24 of the position calculation processing units 5a and 5b temporarily stores the incident timings t and t in the AZD conversion signal storage unit 23 based on the occurrence of the incident timings t and t.
  • Step S3 The scintillator array identifying unit 24 of the position calculation processing units 5a and 5b integrates the digital signals after the AZD conversion by sequentially adding them.
  • T1 2 s X n: n is obtained as a total addition value A up to (total number of additions), and the process proceeds to step S4.
  • Step S4 the scintillator array identification unit 24 of the position calculation processing units 5a and 5b obtains an identification value indicating the intermediate addition value A and the second calculation value A from the intermediate addition value A and the second calculation value A.
  • Step S5 The incident timing correction unit 29 of the position calculation processing units 5a and 5b determines that the scintillator array identified by the scintillator array identification unit 24 has an identification value smaller than the intermediate value K, that is, a light emission pulse. In the incident timing calculation unit 4a, 4b It is determined that correction is made to the calculated incident timing t — At (incident timing correction value).
  • the incident timing correction unit 29 reads the incident timing t stored in the incident timing storage unit 22, and corrects t ⁇ A t for this incident timing t.
  • the incident timing correction unit 29 of the position calculation processing units 5a and 5b determines that the scintillator array force identification value identified by the scintillator array identification unit 24 is larger than the intermediate value, that is, the light emission pulse. It is determined that the incident timing t calculated by the incident timing calculation units 4a and 4b is not corrected. Therefore, the incident tag
  • the coincidence processing unit 6 reads the incident timings t and t stored in the corrected incident timing storage unit every 128 ns, and the incident timings t and t from the two ⁇ -ray detectors 1 and
  • the incident timing is corrected based on the determination result.
  • the detection time difference between the scintillator array 18a and the scintillator array 18b caused by the difference in the decay time of the light emitting panel can be eliminated by correction. Therefore, when scintillators with different emission pulse decay times are used in the shoreline detector 1.
  • the scintillator array identifying unit 24 can identify which scintillator array is the light emitting light emitted by the scintillator.
  • the integration operation that has been conventionally performed by the integrator can be replaced with the addition operation in which the addition means sequentially adds, the number of parts can be reduced and the cost can be reduced.
  • the PET apparatus has been described as an example.
  • the present invention provides a nucleus for performing nuclear medicine diagnosis by simultaneously counting radiation generated from a subject to which a radiopharmaceutical is administered. Any medical device can be applied without being limited to the PET device.
  • the present invention can be applied to an apparatus combining a nuclear medicine diagnostic apparatus and an X-ray CT apparatus, such as a PET-CT equipped with a PET apparatus and an X-ray CT apparatus. it can.
  • the timing window storage unit 32 may store a timing window Tw corresponding to each combination of the plurality of scintillator arrays. Therefore, by using different timing windows Tw depending on the combination of multiple scintillator arrays, it is possible to perform highly accurate coincidence counting, reduce the effects of accidental coincidence counting, scattered clock counting, etc. A high-quality image can be obtained.
  • the timing window storage unit 32 may store a timing window Tw corresponding to each combination of a plurality of scintillators. Therefore, by using different timing windows Tw depending on the combination of multiple scintillators, it is possible to perform highly accurate coincidence, reduce the effects of accidental coincidence and scatter coincidence, reduce noise and achieve high image quality. An image can be obtained.
  • the incident timing correction unit 29 of the position calculation processing units 5a and 5b uses the incident timing t calculated by the incident timing calculation units 4a and 4b as t — At (incident timing).
  • the incident timing correction unit 29 of the position calculation processing units 5a and 5b performs calculation processing with the incident timing t calculated by the incident timing calculation units 4a and 4b as t + At.
  • the incident timing t is not corrected, and the incident timing t + At and the incident timing t are
  • the time difference between the rise times different in the scintillator arrays 18a and 18b is stored in the correction data table 31 as the incident timing correction value.
  • the time difference may be stored in the correction data table 31 as the incident timing correction value. Therefore, it is possible to correct the time difference between the rise times that differ between the scintillators 19, and to further improve the detection sensitivity.
  • the time difference between the rise times different in the scintillator arrays 18a, 18b is stored in advance in the correction data table 31 as the incident timing correction value.
  • the incident timing correction value may be calculated in real time with a simple linear function using the identification value as a variable.
  • the scintillator block 15 is a scintillator block using Lu Y SiO (LYSO) as the scintillator 19a having a short emission pulse decay time on the ⁇ -ray incident side (front stage).
  • LYSO Lu Y SiO
  • GSO Ce-concentrated Gd SiO
  • the scintillator array 18a scintillator 19a and the scintillator array 18b scintillator 19b consist of a Ce concentration of 0.5 mol Gd SiO (GSO) and a Ce concentration. 1. 5mol Gd SiO (GSO), Lu SiO (L
  • X 2-X 5 X 2-X 5 4 3 12 aF and CsF may be selected and used in various combinations.
  • the scintillator block 15 has a force other than two layers (pieces) described as a combination of two layers (pieces) of the scintillator array 18a and the scintillator array 18b. Multiple layers (pieces) may be used.
  • the number of scintillators 19a and 19b provided in each scintillator is described as 8 ⁇ 8, a plurality of other scintillators may be provided! (10)
  • the light receiving element is described as the photomultiplier tube 17, but other light receiving elements such as a photodiode or an avalanche photodiode may be used.

Abstract

A nuclear medical diagnosis device includes: a plurality of scintillator arrays each formed by a plurality of scintillators having different attenuation times of the light emitting pulse in the Ϝ-ray incident depth direction; incidence timing calculation means for calculating the timing of the incidence into the scintillator array; scintillator array identification means for identifying the scintillator into which the Ϝ-ray has come; and an incidence timing correction unit of a position calculation processing unit for judging whether to correct the incidence timing calculated by the incidence timing calculation unit according to the scintillator array identified by the scintillator array identification unit. Even when a scintillator having a different attenuation time of the light emitting pulse is used for the Ϝ-ray detector, the incidence timing can be corrected according to the result of judgment by the incidence timing correction unit so as to improve the detection sensitivity and obtain an accurate tomogram without causing degradation of a reconfigured image.

Description

核医学診断装置  Nuclear medicine diagnostic equipment
技術分野  Technical field
[0001] この発明は、被検体に放射性薬剤が投与され、この被検体の関心部位に蓄積され たポジトロン放射性同位元素（ラジオアイソトープ， RI)から放出された一対の γ線を 同時計測し、関心部位の断層像を得るための核医学診断装置 (ECT装置）に係り、 特に、 γ線を同時計数する技術に関する。  [0001] In the present invention, a radiopharmaceutical is administered to a subject, and a pair of γ-rays emitted from a positron radioisotope (radioisotope, RI) accumulated in a region of interest of the subject is simultaneously measured, and the subject is interested. The present invention relates to a nuclear medicine diagnostic apparatus (ECT apparatus) for obtaining a tomographic image of a region, and in particular, to a technique for simultaneously counting γ rays.
背景技術  Background art
[0002] 上述した核医学診断装置、すなわち ECT(Emission Computed Tomography)装置と して、 PET(Positron Emission Tomography)装置を例に採って説明する。 PET装置は 、被検体の関心部位力 互いにほぼ 180° 方向に放出される 2本の γ線を対向する Ύ線検出器により検出し、これら γ線が同時に検出（同時計数)されたときに被検体 の断層画像を再構成するように構成されている。また、 PET装置で γ線を同時計数 するために用いられる γ線検出器としては、被検体力 放出された γ線が入射して 発光するシンチレータと、このシンチレータでの発光を電気信号に変換する光電子 増倍管とから構成されたものがある。  [0002] As a nuclear medicine diagnosis apparatus, that is, an ECT (Emission Computed Tomography) apparatus, a PET (Positron Emission Tomography) apparatus will be described as an example. The PET device detects the force of the region of interest of the subject by detecting the two gamma rays emitted in a direction of approximately 180 ° from each other with the opposing gutter detectors, and when these gamma rays are simultaneously detected (simultaneously counted). It is configured to reconstruct a tomographic image of the specimen. In addition, the γ-ray detector used to simultaneously count γ-rays with a PET device includes a scintillator that emits light when γ-rays emitted from the subject force enter, and converts the light emitted from the scintillator into an electrical signal. Some are composed of photomultiplier tubes.
[0003] ここで、原理的に視野中心力 離れた位置力 放出される γ線は、 γ線検出器の シンチレータに斜めから入射することが多くなり、正しい位置の検出だけでなぐ誤つ た位置においても検出されることになる。つまり、視野中心力も周辺部に向力つて徐 々に視差誤差が大きくなり、 PET装置で得られる断層画像は不正確なものとなって いる。そこで、シンチレータを γ線入射方向において、発光ノ ルスの減衰時間が異な るシンチレータに分割（光学的に結合）、例えば、シンチレータを γ線入射側に γ線 の減衰時間の短 、シンチレータアレイと、光電子増倍管側に γ線の減衰時間の長 、 シンチレータアレイとに分割し、 γ線が γ線検出器のシンチレータに斜めに入射した 場合でも、放射された γ線の位置を精度よく検出し、より正確な断層画像を得るよう に改善を図っている（例えば、特許文献 1, 2参照)。  [0003] Here, in principle, a positional force that is distant from the central force of the field of view. The emitted γ-rays are often incident on the scintillator of the γ-ray detector from an oblique direction. Will also be detected. In other words, the parallax error gradually increases as the visual field central force is directed toward the periphery, and the tomographic image obtained by the PET apparatus is inaccurate. Therefore, the scintillator is divided (optically coupled) into scintillators with different light emission attenuation times in the direction of γ-ray incidence, for example, the scintillator has a short γ-ray attenuation time on the γ-ray incidence side, The γ-ray decay time is divided into a scintillator array on the photomultiplier tube side, and the position of the emitted γ-ray is detected accurately even when the γ-ray is obliquely incident on the scintillator of the γ-ray detector. Improvements are made to obtain more accurate tomographic images (see, for example, Patent Documents 1 and 2).
特許文献 1 :特開平 6— 337289号公報 (第 2— 3頁、図 1) 特許文献 2：特開 2000 - 56023号公報 (第 2— 3頁、図 1) Patent Document 1: JP-A-6-337289 (Page 2-3, Fig. 1) Patent Document 2: JP 2000-56023 A (Page 2-3, Fig. 1)
発明の開示  Disclosure of the invention
発明が解決しょうとする課題  Problems to be solved by the invention
[0004] し力しながら、従来の核医学診断装置では、次のような問題がある。すなわち、発光 パルスの減衰時間が異なるシンチレータでは、発光パルスの立ち上がり時間も異なる ことが多い。したがって、発光パルスの立ち上がり時間の異なるシンチレータアレイを 用いた γ線検出器により同時計数した場合、対向する γ線検出器間における検出 に時間差が生じることになる。つまり、この時間差が生じることにより、実際には同時に 放射された γ線が、 γ線検出器での検出に基づく同時計数処理において、同時に γ線が放射されたとは認識されず、検出感度を低下させるという問題がある。また、 検出感度の低下を改善するために、同時計数処理での有効な同時計数カウントとす る時間範囲（タイミングウィンド)を広くとり、対向する γ線検出器間における検出に時 間差がある場合でも同時と判別させるようにすると、偶発同時計数や散乱同時計数 などの影響が増大し、再構成画像の劣化を招くという問題がある。  However, the conventional nuclear medicine diagnostic apparatus has the following problems. In other words, scintillators with different emission pulse decay times often have different emission pulse rise times. Therefore, when simultaneous counting is performed by a γ-ray detector using a scintillator array with different rise times of the light emission pulses, a time difference occurs in detection between the opposing γ-ray detectors. In other words, due to this time difference, γ rays that are emitted simultaneously are not recognized as being simultaneously emitted in the coincidence process based on the detection by the γ detector, and the detection sensitivity decreases. There is a problem of making it. In addition, in order to improve the decrease in detection sensitivity, the time range (timing window) for effective coincidence counting in the coincidence counting process is widened, and there is a time difference in detection between opposing γ-ray detectors. Even if it is determined to be simultaneous, there is a problem that the influence of accidental coincidence counting and scattering coincidence counting increases, leading to degradation of the reconstructed image.
[0005] この発明は、このような事情に鑑みてなされたものであって、 γ線検出器に発光パ ルスの減衰時間の異なるシンチレータを用いた場合でも、高感度であり、再構成画像 の劣化を生じさせず、正確な断層画像を得ることができる核医学診断装置を提供す ることを目的とする。  [0005] The present invention has been made in view of such circumstances, and even when a scintillator having a different emission pulse decay time is used for the γ-ray detector, the sensitivity is high, and An object of the present invention is to provide a nuclear medicine diagnostic apparatus capable of obtaining an accurate tomographic image without causing deterioration.
課題を解決するための手段  Means for solving the problem
[0006] この発明は、このような目的を達成するために、次のような構成をとる。 In order to achieve such an object, the present invention has the following configuration.
すなわち、この発明の核医学診断装置は、複数個のシンチレータを 2次元的に密 着配置し、 y線入射深さ方向に発光パルスの減衰時間が異なる複数個のシンチレ ータアレイを光学的に結合したシンチレ一タブロックと、シンチレ一タブロックで発光 した発光パルスを電気信号に変換する受光素子と、受光素子力 出力された電気信 号について、前記シンチレータアレイに入射したタイミングを算出する入射タイミング 算出手段と、受光素子力 出力された電気信号が複数個あるうちのいずれのシンチ レータァレイに入射したかを識別するシンチレータアレイ識別手段と、シンチレータァ レイ識別手段で識別されたシンチレータアレイに応じて、前記入射タイミング算出手 段で算出された入射タイミングの補正を行うか否かの判別を行い、判別の結果に基 づいて、入射タイミングの補正を行う入射タイミング補正手段と、を備えていることを特 徴とするちのである。 That is, in the nuclear medicine diagnosis apparatus of the present invention, a plurality of scintillators are two-dimensionally closely arranged, and a plurality of scintillator arrays having different emission pulse decay times in the y-ray incident depth direction are optically coupled. A scintillator block, a light receiving element that converts a light emission pulse emitted from the scintillator block into an electric signal, and an incident timing calculating means for calculating a timing at which the electric signal output from the light receiving element force is incident on the scintillator array And the light receiving element force. Depending on the scintillator array identifying means for identifying which of the plurality of output electrical signals has entered the scintillator array, and the scintillator array identified by the scintillator array identifying means, the incident Timing calculation hand It is determined whether or not to correct the incident timing calculated in the stage, and incident timing correction means for correcting the incident timing based on the determination result. It is.
[0007] 請求項 1の発明の作用は次のとおりである。まず、被検体力も放出された γ線は、 複数個のシンチレータを 2次元的に密着配置し、 γ線入射深さ方向に発光パルスの 減衰時間が異なる複数個のシンチレータアレイを光学的に結合したシンチレ一タブ ロックに入射する。さらに、シンチレ一タブロックに入射した γ線は、発光パルスの減 衰時間が異なるシンチレータアレイの各シンチレータで発光を行う。さらに、各シンチ レータで発光された発光ノ ルスは、受光素子により電気信号に変換される。次に、入 射タイミング算出手段は、受光素子力も出力された電気信号について、シンチレータ アレイに入射したタイミングを算出する。また、シンチレータアレイ識別手段は、受光 素子から出力された電気信号が複数個あるうちのいずれのシンチレータアレイに入 射したかを識別する。さら〖こ、入射タイミング補正手段は、シンチレータアレイ識別手 段で識別されたシンチレータアレイに応じて、入射タイミング算出手段で算出された 入射タイミングの補正を行うか否かの判別を行い、判別の結果に基づいて、入射タイ ミングの補正を行う。したがって、入射タイミング補正手段で、シンチレータアレイ識別 手段で識別されたシンチレータアレイに応じて、入射タイミング算出手段で算出され た入射タイミングの補正を行うか否かの判別を行い、判別の結果に基づいて、入射タ イミングの補正を行うので、発光パルスの減衰時間の異なるシンチレータアレイを用 いて同時計数した場合においても、この発光パルスの減衰時間の異なることにより生 じるシンチレータアレイ間における検出の時間差を補正により解消することができる。 したがって、検出感度を高め、再構成画像の劣化を生じさせず、正確な断層画像を 得ることができる。  [0007] The operation of the invention of claim 1 is as follows. First, the γ rays from which the subject force was also released are a two-dimensional arrangement of multiple scintillators, and optically combined multiple scintillator arrays with different emission pulse decay times in the direction of the γ-ray incident depth. Incident into the scintillation tab lock. Furthermore, the γ-rays incident on the scintillator block emit light by each scintillator of the scintillator array having a different emission pulse decay time. Furthermore, the light emission light emitted by each scintillator is converted into an electric signal by the light receiving element. Next, the incident timing calculation means calculates the timing at which the light signal is also incident on the scintillator array for the output electrical signal. The scintillator array identifying means identifies which scintillator array among a plurality of electrical signals output from the light receiving elements is incident. Furthermore, the incident timing correction means determines whether or not to correct the incident timing calculated by the incident timing calculation means in accordance with the scintillator array identified by the scintillator array identification means. Based on the above, the incident timing is corrected. Therefore, the incident timing correction means determines whether or not to correct the incident timing calculated by the incident timing calculation means according to the scintillator array identified by the scintillator array identification means, and based on the determination result. Since the incident timing is corrected, even when simultaneous counting is performed using scintillator arrays with different emission pulse decay times, the detection time difference between the scintillator arrays caused by the different decay times of the emission pulses can be reduced. It can be solved by correction. Therefore, it is possible to increase the detection sensitivity and obtain an accurate tomographic image without causing deterioration of the reconstructed image.
[0008] また、請求項 2の発明の核医学診断装置は、受光素子から出力された電気信号で あるアナログ信号をデジタル信号に変換する AZD変 を備え、シンチレ一タァレ ィ識別手段は、 AZD変 で変換されたデジタル信号を順次加算する加算手段と 、加算手段において、シンチレ一タブロックで発光した発光パルスの発光開始時から 発光終了時までの途中である途中時点までのデジタル信号を加算した途中加算値 およびシンチレ一タブロックで発光した発光パルスの発光開始時力 発光終了時ま でのデジタル信号を加算した全加算値から、途中加算値を全加算値で除算した値を 示す識別値を算出する識別値算出手段と、識別値算出手段で算出された各シンチ レータアレイの識別値間の中間値に対して、前記識別値算出手段で算出された識別 値が大き ヽ値か小さ ヽ値かを判別する判別手段とを備えて ヽることを特徴とするもの である。 [0008] In addition, the nuclear medicine diagnosis apparatus of the invention of claim 2 includes an AZD modification that converts an analog signal that is an electric signal output from the light receiving element into a digital signal, and the scintillation array identifying means includes an AZD modification. In the addition means for sequentially adding the digital signals converted in step 1, and in the addition means, in the middle of adding the digital signals from the start of light emission emitted by the scintillator block to the middle of the light emission. Addition value Identification for calculating the identification value indicating the value obtained by dividing the intermediate addition value by the total addition value from the total addition value obtained by adding the digital signals up to the end of light emission of the emission pulse emitted by the scintillator block It is determined whether the identification value calculated by the identification value calculation means is a large threshold value or a small threshold value with respect to an intermediate value between the identification values of the scintillator array calculated by the value calculation means and the identification value calculation means. It is characterized by having a discrimination means.
[0009] この発明の請求項 2の核医学診断装置によれば、 AZD変翻は、受光素子から 出力された電気信号であるアナログ信号をデジタル信号に変換する。次に、シンチレ ータアレイ識別手段の加算手段は、 AZD変換器で変換されたデジタル信号を順次 加算する。また、識別値算出手段は、加算手段での加算により求められたシンチレ一 タブロックで発光した発光パルスの発光開始時力 発光終了時までの途中である途 中時点までのデジタル信号を加算した途中加算値と、シンチレ一タブロックで発光し た発光パルスの発光開始時力 発光終了時までのデジタル信号を加算した全加算 値とから、途中加算値を^ロ算値で除算した値を示す識別値を算出する。さら〖こ、中 間値算出手段は、識別値算出手段で算出された各シンチレータアレイの識別値間の 中間値を求め、判別手段により、中間値算出手段で算出された中間値に対して、識 別値算出手段で算出された識別値が大きい値力 vj、さい値かを判別する。したがって According to the nuclear medicine diagnostic apparatus of claim 2 of the present invention, the AZD conversion converts an analog signal that is an electrical signal output from the light receiving element into a digital signal. Next, the addition means of the scintillator array identification means sequentially adds the digital signals converted by the AZD converter. In addition, the identification value calculation means adds the digital signal up to the midpoint of the light emission start power of the light emission pulse emitted by the scintillator block obtained by addition by the addition means until the end of light emission. Identification that shows the value obtained by dividing the intermediate addition value by the half-calculated value from the addition value and the total addition value obtained by adding the digital signals until the end of light emission of the light emission pulse emitted by the scintillator block Calculate the value. Furthermore, the intermediate value calculation means obtains an intermediate value between the identification values of each scintillator array calculated by the identification value calculation means, and the determination means determines the intermediate value calculated by the intermediate value calculation means. It is determined whether the discriminant value calculated by the discriminant value calculation means is a large value vj or a small value. Therefore
、シンチレータアレイ識別手段の加算手段で順次加算することに基づいて、識別値 算出手段で算出された識別値が大きい値力 vj、さい値かを判別することができる。つま り、シンチレータアレイ識別手段は、シンチレータで発光した発光パルスが何れのシ ンチレータァレイであるかを識別することができる。また、従来積分器で行われていた 積分動作を加算手段で順次加算する加算動作に置き換えることができるので、部品 点数の削減とコストダウンを図ることができる。 Based on the sequential addition by the addition means of the scintillator array identification means, it is possible to determine whether the identification value calculated by the identification value calculation means is a large value force vj or small value. In other words, the scintillator array identifying means can identify which scintillator array is the light emission pulse emitted by the scintillator. In addition, since the integration operation that has been conventionally performed by the integrator can be replaced with an addition operation in which the addition means sequentially adds, the number of parts can be reduced and the cost can be reduced.
[0010] また、請求項 3の発明の核医学診断装置は、入射タイミング補正手段で補正した入 射タイミングおよび入射タイミング補正手段で入射タイミングを補正しないと判別され た入射タイミングを用いて同時計数を行う同時計数手段と、同時計数手段で同時計 数とする所定の範囲を示すタイミングウィンドを複数個のシンチレータアレイそれぞれ の組み合わせに対応するタイミングウィンドとして記憶するタイミングウィンド記憶手段 と、を備えていることを特徴とするものである。 [0010] Further, the nuclear medicine diagnosis apparatus of the invention of claim 3 performs simultaneous counting using the incident timing corrected by the incident timing correcting means and the incident timing determined not to correct the incident timing by the incident timing correcting means. And a timing window storage means for storing a timing window indicating a predetermined range of the same clock number by the simultaneous counting means as a timing window corresponding to each combination of a plurality of scintillator arrays It is characterized by having these.
[0011] この発明の請求項 3の核医学診断装置によれば、同時計数手段は、入射タイミング 補正手段で補正した入射タイミングおよび入射タイミング補正手段で入射タイミングを 補正しないと判別された入射タイミングを用いて同時計数を行う。また、同時計数は、 同時計数が同時と判定される所定の範囲を示すタイミングウィンドを用いて行われ、 ここでは、タイミングウィンド記憶手段に記憶された複数個のシンチレータアレイそれ ぞれの組み合わせに対応するタイミングウィンドを用いて同時計数される。したがって 、複数個のシンチレータアレイそれぞれの組み合わせにより異なるタイミングウィンド を用いることで、精度の高い同時計数を行うことができ、偶発同時計数や散乱同時計 数などの影響を減らし、ノイズの少な 、高画質な画像を得ることができる。  According to the nuclear medicine diagnostic apparatus of claim 3 of the present invention, the coincidence counting unit calculates the incident timing corrected by the incident timing correcting unit and the incident timing determined not to correct the incident timing by the incident timing correcting unit. To perform coincidence. In addition, the coincidence counting is performed using a timing window indicating a predetermined range in which the coincidence counting is determined to be simultaneous, and here, it corresponds to a combination of a plurality of scintillator arrays stored in the timing window storage means. The timing window is used for simultaneous counting. Therefore, by using different timing windows depending on the combination of each of the multiple scintillator arrays, it is possible to perform highly accurate coincidence, reduce the effects of accidental coincidence and scattered clocks, reduce noise, and achieve high image quality. Can be obtained.
[0012] また、請求項 4の発明の核医学診断装置は、入射タイミング補正手段で補正した入 射タイミングおよび入射タイミング補正手段で入射タイミングを補正しないと判別され た入射タイミングを用いて同時計数を行う同時計数手段と、同時計数手段で同時計 数とする所定の範囲を示すタイミングウィンドを複数個のシンチレータそれぞれの組 み合わせに対応するタイミングウィンドとして記憶するタイミングウィンド記憶手段と、 を備えて ヽることを特徴とするものである。  [0012] Further, the nuclear medicine diagnosis apparatus of the invention of claim 4 performs simultaneous counting using the incident timing corrected by the incident timing correcting means and the incident timing determined not to correct the incident timing by the incident timing correcting means. And a timing window storage means for storing a timing window indicating a predetermined range of the same clock number by the simultaneous counting means as a timing window corresponding to each combination of a plurality of scintillators. It is characterized by that.
[0013] この発明の請求項 4の核医学診断装置によれば、同時計数手段は、入射タイミング 補正手段で補正した入射タイミングおよび入射タイミング補正手段で入射タイミングを 補正しないと判別された入射タイミングを用いて同時計数を行う。また、同時計数は、 同時計数が同時と判定される所定の範囲を示すタイミングウィンドを用いて行われ、 ここでは、タイミングウィンド記憶手段に記憶された複数個のシンチレータそれぞれの 組み合わせに対応するタイミングウィンドを用いて同時計数される。したがって、複数 個のシンチレータそれぞれの組み合わせにより異なるタイミングウィンドを用いること で、精度の高い同時計数を行うことができ、偶発同時計数や散乱同時計数などの影 響を減らし、ノイズの少な 、高画質な画像を得ることができる。  According to the nuclear medicine diagnostic apparatus of claim 4 of the present invention, the coincidence counting unit calculates the incident timing corrected by the incident timing correcting unit and the incident timing determined not to correct the incident timing by the incident timing correcting unit. To perform coincidence. In addition, the coincidence counting is performed using a timing window indicating a predetermined range in which the coincidence counting is determined to be simultaneous. Here, a timing window corresponding to each combination of a plurality of scintillators stored in the timing window storage means. Are simultaneously counted. Therefore, by using different timing windows depending on the combination of multiple scintillators, it is possible to perform highly accurate coincidence counting, reduce the effects of accidental coincidence counting and scattering coincidence counting, reduce noise and achieve high image quality. An image can be obtained.
[0014] また、請求項 5の発明の核医学診断装置は、シンチレ一タブロックと受光素子とを 光学的に結合するライトガイドを備えていることを特徴とするものである。  [0014] Further, the nuclear medicine diagnosis apparatus of the invention of claim 5 is characterized by comprising a light guide for optically coupling the scintillator block and the light receiving element.
[0015] この発明の請求項 5の核医学診断装置によれば、シンチレ一タブロックと受光素子 とを光学的に結合するライトガイドを備えている。したがって、ライトガイドにより、シン チレ一タブロックからの光を受光素子に適切に導くことができる。 According to the nuclear medicine diagnostic apparatus of claim 5 of the present invention, the scintillator block and the light receiving element And a light guide that optically couples the two. Therefore, the light guide can appropriately guide the light from the scintillator block to the light receiving element.
[0016] また、請求項 6の発明の核医学診断装置は、複数個のシンチレータアレイは、 Ce濃 度 0. 5molの Gd SiO (GSO) , Ce濃度 1. 5molの Gd SiO (GSO) , Lu SiO (LS [0016] Further, in the nuclear medicine diagnostic apparatus of the invention of claim 6, the plurality of scintillator arrays are made of Gd SiO (GSO) having a Ce concentration of 0.5 mol, Gd SiO (GSO), Lu having a Ce concentration of 1.5 mol. SiO (LS
2 5 2 5 2 5 2 5 2 5 2 5
O) , Lu Gd SiO (LGSO) , Lu Y SiO (LYSO) , Bi Ge 0 (BGO) , Nal, BaO), Lu Gd SiO (LGSO), Lu Y SiO (LYSO), Bi Ge 0 (BGO), Nal, Ba
X 2-X 5 X 2-X 5 4 3 12 X 2-X 5 X 2-X 5 4 3 12
F , CsFのいずれかのシンチレータにより構成されていることを特徴とするものである It is characterized by being composed of either F or CsF scintillator
2 2
[0017] この発明の請求項 6の核医学診断装置によれば、複数個のシンチレータアレイは、 Ce濃度 0. 5molの Gd SiO (GSO) , Ce濃度 1. 5molの Gd SiO (GSO) , Lu SiO According to the nuclear medicine diagnostic apparatus of claim 6 of the present invention, the plurality of scintillator arrays are made of Gd SiO (GSO) having a Ce concentration of 0.5 mol, Gd SiO (GSO), Lu having a Ce concentration of 1.5 mol. SiO
2 5 2 5 2 2 5 2 5 2
(LSO) , Lu Gd SiO (LGSO) , Lu Y SiO (LYSO) , Bi Ge 0 (BGO) , Nal(LSO), Lu Gd SiO (LGSO), Lu Y SiO (LYSO), Bi Ge 0 (BGO), Nal
5 X 2-X 5 X 2-X 5 4 3 12 5 X 2-X 5 X 2-X 5 4 3 12
， BaF， CsFにより構成されている。したがって、複数個のシンチレータアレイを構成 , BaF, CsF. Therefore, configure multiple scintillator arrays
2 2
するシンチレータを種々選択することができ、高価なシンチレータだけでなぐ安価な シンチレータも使用することができ、コストの削減を行うことができる。  It is possible to select various scintillators to be used, and it is possible to use not only expensive scintillators but also inexpensive scintillators, thereby reducing costs.
[0018] また、請求項 7の発明の核医学診断装置は、受光素子は光電子増倍管であること を特徴とするものである。 [0018] Further, the nuclear medicine diagnosis apparatus of the invention of claim 7 is characterized in that the light receiving element is a photomultiplier tube.
[0019] この発明の請求項 7の核医学診断装置によれば、受光素子は光電子増倍管である ので、シンチレ一タブロックからの光を適切に電気信号に変換することができる。 According to the nuclear medicine diagnostic apparatus of claim 7 of the present invention, since the light receiving element is a photomultiplier tube, the light from the scintillator block can be appropriately converted into an electric signal.
[0020] また、請求項 8の発明の核医学診断装置は、受光素子は受光素子はフォトダイォ ードであることを特徴とするものである。 [0020] Further, the nuclear medicine diagnosis apparatus of the invention of claim 8 is characterized in that the light receiving element is a photodiode.
[0021] この発明の請求項 8の核医学診断装置によれば、受光素子は受光素子はフォトダ ィオードであるので、シンチレ一タブロックからの光を適切に電気信号に変換すること ができる。 [0021] According to the nuclear medicine diagnosis apparatus of claim 8 of the present invention, since the light receiving element is a photodiode, the light from the scintillator block can be appropriately converted into an electric signal.
[0022] また、請求項 9の発明の核医学診断装置は、受光素子はァバランシ フォトダイォ ードであることを特徴とするものである。  [0022] Further, the nuclear medicine diagnosis apparatus of the invention of claim 9 is characterized in that the light receiving element is an avalanche photodiode.
[0023] この発明の請求項 9の核医学診断装置によれば、受光素子はアバランシェフオトダ ィオードであるので、シンチレ一タブロックからの光を適切に電気信号に変換すること ができる。 According to the nuclear medicine diagnostic apparatus of claim 9 of the present invention, since the light receiving element is an avalanche photodiode, it is possible to appropriately convert the light from the scintillator block into an electric signal.
発明の効果 [0024] この発明に係る核医学診断装置によれば、入射タイミング補正手段で、シンチレ一 タアレイ識別手段で識別されたシンチレータアレイに応じて、入射タイミング算出手段 で算出された入射タイミングの補正を行うか否かの判別を行い、判別の結果に基づ いて、入射タイミングの補正を行うので、発光パルスの減衰時間の異なるシンチレ一 タアレイを用いて同時計数した場合においても、この発光パルスの減衰時間の異な ることにより生じるシンチレータアレイ間における検出の時間差を補正により解消する ことができる。したがって、検出感度を高め、再構成画像の劣化を生じさせず、正確 な断層画像を得ることができる。 The invention's effect [0024] According to the nuclear medicine diagnosis apparatus of the present invention, the incident timing correction unit corrects the incident timing calculated by the incident timing calculation unit according to the scintillator array identified by the scintillator array identification unit. Since the incident timing is corrected based on the result of the determination, the decay time of the light emission pulse can be obtained even when simultaneous counting is performed using a scintillator array having a different decay time of the light emission pulse. The time difference of detection between the scintillator arrays caused by the difference between the two can be eliminated by correction. Therefore, it is possible to increase the detection sensitivity and obtain an accurate tomographic image without causing deterioration of the reconstructed image.
図面の簡単な説明  Brief Description of Drawings
[0025] [図 1]PET装置の全体構成を示すブロック図である。  FIG. 1 is a block diagram showing an overall configuration of a PET apparatus.
[図 2]FPGAの構成を示すブロック図である。  FIG. 2 is a block diagram showing a configuration of an FPGA.
[図 3] y線検出器の構成を示した斜視図である。  FIG. 3 is a perspective view showing a configuration of a y-ray detector.
[図 4]増幅回路から出力された各シンチレータアレイの発光パルスを示すグラフであ る。  FIG. 4 is a graph showing light emission pulses of each scintillator array output from the amplifier circuit.
[図 5] (a) , (b)は、各シンチレータアレイに入射した γ線のタイミングを示す図である  [FIG. 5] (a) and (b) are diagrams showing the timing of γ rays incident on each scintillator array.
[図 6]発光パルスの発光開始時力も発光終了時までの加算値を示すグラフである。 FIG. 6 is a graph showing the addition power until the end of light emission as well as the light emission start force of the light emission pulse.
[図 7]タイミングウィンドを説明するためのグラフである。  FIG. 7 is a graph for explaining a timing window.
[図 8]シンチレータアレイの減衰時間が異なることによる検出時間の差を補正していな V、場合のタイミングスペクトルを示すグラフである。  FIG. 8 is a graph showing the timing spectrum when V is not corrected for the difference in detection time due to different decay times of the scintillator array.
[図 9]シンチレータアレイの減衰時間が異なることによる検出時間の差を補正した場 合のタイミングスペクトルを示すグラフである。  FIG. 9 is a graph showing a timing spectrum when a difference in detection time due to different decay times of the scintillator array is corrected.
[図 10]増幅回路力も出力された後の信号処理を示すフローチャートである。  FIG. 10 is a flowchart showing signal processing after the amplification circuit power is also output.
符号の説明  Explanation of symbols
[0026] 3a, 3b … AZD変^^ [0026] 3a, 3b… AZD strange ^^
4a, 4b … 入射タイミング算出部 (入射タイミング算出手段）  4a, 4b ... Incident timing calculator (incident timing calculator)
6 … 同時計数処理部（同時計数手段）  6 ... Simultaneous counting processing unit (simultaneous counting means)
15 … シンチレ一タブロック 16 … ライトガイド 15… Scintillator block 16… Light guide
17 … 光電子増倍管 (受光素子）  17… Photomultiplier tube (light receiving element)
18a, 18b … シンチレータアレイ  18a, 18b… scintillator array
19a, 19b … シンチレータ  19a, 19b… scintillator
24 … シンチレータアレイ識別部 (シンチレータアレイ識別手段）  24… Scintillator array identification unit (scintillator array identification means)
25 … 加算部 (加算手段）  25… Adder (addition means)
26 … 識別値算出部 (識別値算出手段）  26… Identification value calculation unit (identification value calculation means)
28 … 判別部 (判別手段）  28… Discrimination part (discrimination means)
29 … 入射タイミング補正部 (入射タイミング補正手段）  29… Incident timing correction unit (incident timing correction means)
32 … タイミングウィンド記憶部（タイミングウィンド記憶手段）  32… Timing window storage (timing window storage means)
Tw … タイミングウィンド  Tw… Timing window
発明を実施するための最良の形態  BEST MODE FOR CARRYING OUT THE INVENTION
[0027] γ線検出器に発光パルスの減衰時間の異なるシンチレータを用いた場合でも、高 感度であり、再構成画像の劣化を生じさせず、正確な断層画像を得るという目的を実 現した。 [0027] Even when a scintillator with a different emission pulse decay time is used for the γ-ray detector, the purpose of obtaining an accurate tomographic image without causing deterioration of the reconstructed image is realized.
実施例  Example
[0028] PET(Positron Emission Tomography)装置を図面に基づいて詳細に説明する。図 1 は、 PET装置の全体構成を示すブロック図である。図 2は、 FPGA7の構成を示すブ ロック図である。なお、本実施例では、核医学診断装置として、 PET装置を例に採つ て説明する。  [0028] A PET (Positron Emission Tomography) apparatus will be described in detail with reference to the drawings. Fig. 1 is a block diagram showing the overall configuration of the PET apparatus. FIG. 2 is a block diagram showing the configuration of FPGA7. In the present embodiment, a PET apparatus will be described as an example of a nuclear medicine diagnosis apparatus.
[0029] PET装置の全体の構成について図 1を用いて説明する。図 1に示すように、 PET 装置は、被検体 Mに放射性薬剤が投与され、この被検体 Mの関心部位に蓄積され たポジトロン放射性同位元素（ラジオアイソトープ， RI)から放出された γ線を入射し て、光を生じ、この光を電気信号に変換して出力する γ線検出器 1を備えている。こ の Ύ線検出器 1は、被検体 Μの体軸回り、例えば、直径 700mm程度の大きさのリン グ状に隙間無く配置（図 1では、 2つの γ線検出器 1のみ図示）されている。したがつ て、被検体の関心部位力も互いにほぼ 180° 方向に放出される 2本の γ線は、対向 する Ύ線検出器 1により検出され、電気信号に変換されて出力される。 [0030] また、 γ線検出器 1から出力された電気信号を増幅する増幅回路 2a, 2bと、この増 幅回路 2a, 2bで増幅されたアナログ信号をデジタル信号に変換する AZD変換器 3 a, 3bと、増幅回路 2a, 2bで増幅された電気信号を入力し、 γ線検出器 1で検出され た γ線が入射した入射タイミングを算出する入射タイミング算出部 4a, 4bと、この AZ D変換器 3a, 3bで変換されたデジタル信号に基づいて、被検体 Mから放出された γ 線が入射した γ線検出器 1の位置を演算する位置演算処理部 5と、位置演算処理部 5と入射タイミング算出部 4a, 4bとからの情報に基づいて、これら 2つの γ線検出器 1 において、 γ線が同時に入射したことを検出（同時計数)する処理を行う同時計数処 理部 6とを備えている。なお、また、図 2に示すように、位置演算処理部 5と同時計数 処理部 6とは、同じ FPGA (Field Programmable Gate Array) 7と呼ばれるプログラミン グ可能な LSI (大規模集積回路）に備えられている。なお、 FPGA7は CPU8, ROM 9, RAM10などの機能を備えたものであり、位置演算処理部 5と同時計数処理部 6と は、 FPGA7の CPU8の一機能である。さらに、図 1に示すように、同時計数処理部 6 で γ線が同時に検出（同時計数)されたと判別された場合には、被検体の断層画像 を再構成する再構成部 11とを備えている。 [0029] The overall configuration of the PET apparatus will be described with reference to FIG. As shown in Fig. 1, the PET device receives a radiopharmaceutical to subject M and enters γ-rays emitted from positron radioisotopes (radioisotopes, RI) accumulated in the region of interest of subject M. Thus, a γ-ray detector 1 that generates light, converts the light into an electrical signal, and outputs the signal is provided. The X-ray detector 1 is arranged around the body axis of the subject X, for example, in a ring shape with a diameter of about 700 mm without any gap (in FIG. 1, only two γ-ray detectors 1 are shown). Yes. Therefore, the two gamma rays emitted from the subject's region of interest in the direction of approximately 180 ° are detected by the opposing shoreline detector 1, converted into an electrical signal, and output. [0030] In addition, amplifier circuits 2a and 2b that amplify the electrical signal output from the γ-ray detector 1, and an AZD converter 3a that converts the analog signal amplified by the amplifier circuits 2a and 2b into a digital signal. , 3b and the electric signals amplified by the amplifier circuits 2a, 2b are input, and the incident timing calculation units 4a, 4b for calculating the incident timing at which the γ-ray detected by the γ-ray detector 1 is incident, and this AZ D Based on the digital signals converted by the converters 3a and 3b, a position calculation processing unit 5 for calculating the position of the γ-ray detector 1 on which the γ rays emitted from the subject M are incident, and a position calculation processing unit 5 Based on the information from the incident timing calculation units 4a and 4b, the two γ-ray detectors 1 include a coincidence processing unit 6 that performs processing for detecting (simultaneous counting) that γ rays are incident simultaneously. I have. In addition, as shown in FIG. 2, the position calculation processing unit 5 and the coincidence counting processing unit 6 are provided in a programmable LSI (Large Scale Integrated Circuit) called an FPGA (Field Programmable Gate Array) 7. ing. The FPGA 7 has functions such as CPU 8, ROM 9, and RAM 10, and the position calculation processing unit 5 and the coincidence processing unit 6 are functions of the CPU 8 of the FPGA 7. Further, as shown in FIG. 1, when the coincidence processing unit 6 determines that γ rays are simultaneously detected (simultaneous counting), the reconstruction unit 11 reconstructs a tomographic image of the subject. Yes.
[0031] その他にも、本実施例装置は、コントローラ 12とモニタ 13と入力部 14などを備えて いる。以下、この実施例装置の各部の構成を具体的に説明する。  In addition, the apparatus of this embodiment includes a controller 12, a monitor 13, an input unit 14, and the like. Hereinafter, the structure of each part of this Example apparatus is demonstrated concretely.
[0032] γ線検出器 1の構成について、図 3を用いて説明する。図 3は γ線検出器 1の構成 を示した斜視図である。図 3に示すように、 γ線検出器 1は、シンチレータ 19を γ線 入射深さ方向にも分割して配置、つまり、シンチレータを 3次元的に配置した DOI (D epth Of lnteraction)検出器である。例えば、この DOI検出器は、シンチレ一タブロッ ク 15とライトガイド 16と光電子増倍管 (PMT) 17とから構成されるものである。  The configuration of the γ-ray detector 1 will be described with reference to FIG. FIG. 3 is a perspective view showing the configuration of the γ-ray detector 1. As shown in FIG. 3, the γ-ray detector 1 is a DOI (Depth Of lnteraction) detector in which the scintillator 19 is divided and arranged in the γ-ray incident depth direction, that is, the scintillator is arranged three-dimensionally. is there. For example, this DOI detector is composed of a scintillator block 15, a light guide 16, and a photomultiplier tube (PMT) 17.
[0033] シンチレ一タブロック 15は、 γ線入射深さ方向（Z方向）に発光パルスの減衰時間 が異なる 2個のシンチレータアレイ 18aとシンチレータアレイ 18bを光学的に結合した ものであり、シンチレータアレイ 18aは複数個のシンチレータ 19aを、シンチレータァ レイ 18bはシンチレータ 19bをそれぞれ 2次元的に密着配置したものである。具体的 に、シンチレ一タブロック 15は、 γ線入射側（前段）に発光パルスの減衰時間が短い シンチレータ 19a (例えば、 Lu Y SiO (LYSO) )を用いたシンチレータアレイ 18a と、ライトガイド 16側（後段）に発光ノ ルスの減衰時間が長いシンチレータ 19b (例え ば、 Ce濃度 0. 5molの Gd SiO (GSO) )を用いたシンチレータアレイ 18bとの 2段（2 [0033] The scintillator block 15 is an optically coupled two scintillator array 18a and scintillator array 18b having different emission pulse decay times in the γ-ray incident depth direction (Z direction). 18a is a plurality of scintillators 19a, and scintillator array 18b is a two-dimensional arrangement of scintillators 19b. Specifically, the scintillator block 15 includes a scintillator array 18a using a scintillator 19a (for example, Lu Y SiO (LYSO)) with a short decay time of the emission pulse on the γ-ray incident side (front stage). And a scintillator array 18b that uses a scintillator 19b (for example, Gd SiO (GSO) with a Ce concentration of 0.5 mol) on the light guide 16 side (rear stage) with a long decay time of the emission light.
2 5  twenty five
個）重ねられたものである。これらシンチレータ 19a, 19bでは、被検体 Mから放出さ れた γ線を入射して発光する。この時、シンチレータ 19a, 19bでは、発光パルスの 減衰時間が異なることから、立ち上がり時間も異なり、減衰時間が長いほど立ち上が り時間が遅くなる。つまり、シンチレータ 19a, 19b間での検出時間に差が生じること になる。また、 2個のシンチレータアレイ 18a, 18bは、それぞれ 8 X 8本 (X方向， Y方 向）のチップ状のシンチレータ 19a, 19bで構成され、また、シンチレータアレイ 18a, 18b内で隣り合うシンチレータ 19a間およびシンチレータ 19b間には γ線が入射して 生じた光を X方向と Υ方向に比例配分させるための光反射材ゃ光透過材および光学 接着剤が場所により挿入または充填されている。  Pieces) In these scintillators 19a and 19b, γ-rays emitted from the subject M are incident to emit light. At this time, in the scintillators 19a and 19b, since the decay times of the light emission pulses are different, the rise times are also different. The longer the decay time, the slower the rise time. That is, a difference occurs in the detection time between the scintillators 19a and 19b. The two scintillator arrays 18a and 18b are each composed of 8 × 8 chip-shaped scintillators 19a and 19b (X direction and Y direction), and adjacent scintillators 19a and 18b in the scintillator arrays 18a and 18b. Between the space and the scintillator 19b, a light reflecting material, a light transmissive material, and an optical adhesive for proportionally distributing the light generated by the incidence of γ rays in the X and Υ directions are inserted or filled depending on the location.
[0034] ライトガイド 16は、シンチレ一タブロック 15のシンチレータ 19a, 19bで生じた光を光 電子増倍管 17に導くものであり、シンチレ一タブロック 15と光電子増倍管 17との間 に挿入され、それぞれ光学接着剤で互いに光学的に結合されて 、る。  The light guide 16 guides light generated by the scintillators 19 a and 19 b of the scintillator block 15 to the photomultiplier tube 17, and is interposed between the scintillator block 15 and the photomultiplier tube 17. Inserted and each optically bonded with an optical adhesive.
[0035] 光電子増倍管 17は、例えば、光電変換膜が 4面 (チャンネル）内蔵したものであり、 シンチレータ 19a, 19bで生じた光は 4面の PMT光電変換膜に入射され、電子増幅 された後、最終的に電気信号 (アナログ信号）に変換されて出力される。したがって、 この光電子増倍管 17での出力が γ線検出器 1の出力となる。なお、上述した光電子 増倍管 17は、受光素子に相当する。  [0035] The photomultiplier tube 17 has, for example, four photoelectric conversion films (channels) built in, and the light generated by the scintillators 19a and 19b is incident on the four-surface PMT photoelectric conversion films and is electronically amplified. After that, it is finally converted into an electrical signal (analog signal) and output. Therefore, the output from the photomultiplier tube 17 becomes the output of the γ-ray detector 1. The photomultiplier tube 17 described above corresponds to a light receiving element.
[0036] 入射タイミング算出部 4a, 4bについて、図 1,図 4〜図 5 (b)を用いて説明する。図 4 は、増幅回路 2aまたは増幅回路 2bから出力されたシンチレータアレイ 18a, 18bの 発光パルスを示すグラフである。図 5 (a) ,図 5 (b)は、シンチレータアレイ 18a, 18b に入射した γ線のタイミングを示す図である。なお、図 4,図 5 (a)に示す (Α)の曲線 は、発光パルスの減衰時間が短いシンチレータアレイ 18aに入射したものを示し、 (B )の曲線は、発光ノ ルスの減衰時間が長いシンチレータアレイ 18bに入射したものを 示す。図 1に示すように、入射タイミング算出部 4a, 4bには、 γ線検出器 1から出力さ れた電気信号が増幅回路 2a, 2bを介して入力され、この電気信号に基づいて、 γ線 検出器 1のシンチレータアレイ 18a, 18bに入射した γ線の入射タイミングを算出する ものである。具体的には、入射タイミング算出部 4a, 4bは、 ARC (Amplitude and Rise -time Compensation) 20と呼ばれる波高および立ち上がり時間補償回路とタイミング 発生回路 21とからなる。 The incident timing calculation units 4a and 4b will be described with reference to FIGS. 1, 4 to 5B. FIG. 4 is a graph showing light emission pulses of the scintillator arrays 18a and 18b output from the amplifier circuit 2a or the amplifier circuit 2b. FIGS. 5 (a) and 5 (b) are diagrams showing the timing of γ rays incident on the scintillator arrays 18a and 18b. The curve (Α) shown in Figs. 4 and 5 (a) shows the light incident on the scintillator array 18a with a short decay time of the emission pulse, and the curve (B) shows the decay time of the emission pulse. Shown incident on long scintillator array 18b. As shown in FIG. 1, the electrical signals output from the γ-ray detector 1 are input to the incident timing calculation units 4a and 4b via the amplifier circuits 2a and 2b, and γ-rays are generated based on the electrical signals. Calculate the incident timing of gamma rays incident on scintillator arrays 18a and 18b of detector 1. Is. Specifically, the incident timing calculation units 4 a and 4 b include a wave height and rise time compensation circuit called ARC (Amplitude and Rise-time Compensation) 20 and a timing generation circuit 21.
[0037] ARC20では、 γ線検出器 1のシンチレータアレイ 18a, 18bに入射した γ線に基 づいて、例えば、図 4に示されるような、増幅回路 2a, 2bから出力された発光パルス の減衰時間が異なるアナログ信号が入力される。さらに、 ARC20は、これらアナログ 信号のそれぞれについて、シンチレータアレイ 18a, 18bに入射した γ線の入射タイ ミングを算出するための波形整形処理を行う。具体的に、この波形整形処理は、増幅 回路 2a, 2bから得られた信号を遅延させたものと、増幅回路 2a, 2bから得られた信 号の電圧値を反転し低くさせたものとを加算演算することで得ることができ、図 5 (a) に示すような電圧波形に整形される。ここで、電圧が 0の時点（ゼロクロスポイント）で ある t , t がシンチレータアレイ 18a, 18bに入射した γ線の入射タイミングを示すも[0037] In the ARC 20, based on the γ-rays incident on the scintillator arrays 18a and 18b of the γ-ray detector 1, for example, attenuation of the light emission pulses output from the amplifier circuits 2a and 2b as shown in FIG. Analog signals with different times are input. Further, the ARC 20 performs a waveform shaping process for calculating the incident timing of the γ rays incident on the scintillator arrays 18a and 18b for each of these analog signals. Specifically, in this waveform shaping process, the signal obtained from the amplification circuits 2a and 2b is delayed and the voltage value of the signal obtained from the amplification circuits 2a and 2b is inverted and lowered. It can be obtained by addition operation, and is shaped into a voltage waveform as shown in Fig. 5 (a). Here, when the voltage is 0 (zero cross point), t and t indicate the incident timing of the γ-rays incident on the scintillator arrays 18a and 18b.
SF SR SF SR
のである。さらに、タイミング発生回路 21は、図 5 (b)に示されるような ARC20で算出 された入射タイミングを示す信号をデジタル信号に変換して、 FPGA7の RAM10の 一機能である入射タイミング記憶部 22に一時的に記憶される構成となって 、る。なお 、上述した入射タイミング算出部 4a, 4bは、入射タイミング算出手段に相当する。  It is. Further, the timing generation circuit 21 converts the signal indicating the incident timing calculated by the ARC 20 as shown in FIG. 5 (b) into a digital signal, and stores it in the incident timing storage unit 22 which is one function of the RAM 10 of the FPGA 7. The configuration is temporarily stored. The incident timing calculation units 4a and 4b described above correspond to incident timing calculation means.
[0038] 位置演算処理部 5について、図 1を用いて説明する。まず、図 1に示すように、 γ線 検出器 1から出力された電気信号が増幅回路 2a, 2bを介して入力され、さらに、この 増幅回路 2a, 2bから入力されたアナログ信号を常時、 AZD変換する AZD変翻 3a, 3bで変換されたデジタル信号力FPGA7の RAM10の一機能である AZD変換 信号記憶部 23に一時的に記憶される。この AZD変換信号記憶部 23に記憶された デジタル信号に基づいて、位置演算処理部 5a, 5bは、被検体 Mから放出された γ 線が入射した γ線検出器 1のシンチレータ 19a, 19bの位置を特定するための演算 処理を行う。ここで、 γ線検出器 1のシンチレータ 19a, 19bの X方向と Y方向（同じシ ンチレータ 18a, 18b内）については、 γ線検出器 1のシンチレータ 19a, 19bの後段 にある 4入力 PMTへの光の配分に基づ 、た電圧値力 位置演算される（アンガー方 式)。 The position calculation processing unit 5 will be described with reference to FIG. First, as shown in FIG. 1, the electrical signal output from the γ-ray detector 1 is input via the amplifier circuits 2a and 2b, and the analog signal input from the amplifier circuits 2a and 2b is always converted to AZD. AZD conversion to be converted The digital signal power converted by the 3a, 3b is temporarily stored in the AZD conversion signal storage unit 23 which is a function of the RAM 7 of the FPGA 7. Based on the digital signal stored in the AZD conversion signal storage unit 23, the position calculation processing units 5a and 5b detect the positions of the scintillators 19a and 19b of the γ-ray detector 1 on which the γ-rays emitted from the subject M are incident. Performs arithmetic processing to identify Here, with respect to the X direction and Y direction (within the same scintillators 18a and 18b) of the scintillators 19a and 19b of the γ-ray detector 1, they are connected to the 4-input PMT at the subsequent stage of the scintillators 19a and 19b of the γ-ray detector 1. Based on the light distribution, the voltage value force position is calculated (Anger method).
[0039] また、位置演算処理部 5a, 5bは、図 2に示すように、 γ線検出器 1で検出された γ 線力 2個あるうちのいずれのシンチレータアレイ 18a, 18bへの入射に基づいて検 出されたものであるかを識別するシンチレータアレイ識別部 24を備えて 、る。言 ヽ換 えれば、このシンチレータアレイ 18aまたはシンチレータアレイ 18bが識別されること により、 Ί線検出器 1のシンチレータ 19aまたはシンチレータ 19bの何れであるかを示 す Z方向にっ 、ての位置演算されることになる。 In addition, as shown in FIG. 2, the position calculation processing units 5a and 5b are configured to detect γ detected by the γ-ray detector 1. A scintillator array identifying unit 24 for identifying which of the two line forces is detected based on the incident on the scintillator arrays 18a and 18b is provided. In other words, when this scintillator array 18a or scintillator array 18b is identified, the position is calculated in the Z direction indicating whether it is the scintillator 19a or scintillator 19b of the winding detector 1. Will be.
シンチレータアレイ識別部 24について、図 6を用いて説明する。図 6は、発光パル スの発光開始時から発光終了時 Tまでの加算値を示すグラフである。なお、図 6に  The scintillator array identification unit 24 will be described with reference to FIG. FIG. 6 is a graph showing the added value from the start of light emission to the end T of light emission. Figure 6 shows
2  2
示す (A)の曲線は、発光パルスの減衰時間が短 、シンチレータ 19a (シンチレータァ レイ 18a)に入射したものを示し、（B)の曲線は、発光パルスの減衰時間が長いシン チレータ 19b (シンチレータアレイ 18b)に入射したものを示す。 A/D変換器 3a, 3b で変換されたデジタル信号を順次加算する加算部 25と、この加算部 25において、シ ンチレータ 19a, 19bで発光した発光ノ ルスの発光開始時力も発光終了時 Tまでの The curve in (A) shows that the emission pulse has a short decay time and is incident on the scintillator 19a (scintillator array 18a). The curve in (B) shows the scintillator 19b (scintillator 19b in which the emission pulse has a long decay time. Shown incident on array 18b). An adder 25 that sequentially adds the digital signals converted by the A / D converters 3a and 3b, and in this adder 25, the light emission starting power of the light emitting light emitted by the scintillators 19a and 19b is also until the light emission end T. of
2 途中である途中時点 Tまでのデジタル信号を加算した途中加算値 A と、シンチレ  2 A halfway addition value A obtained by adding the digital signals up to midway point T and a scintillation
1 T1 一 タ 19a, 19bで発光した発光パルスの発光開始時力 発光終了時 T2までのデジタル 信号を加算した全加算値 A とから、途中加算値 A 加算値 A (途中加算値 A  1 T1 counter Light emission start force of light emission pulses emitted by 19a, 19b At the end of light emission From the total addition value A obtained by adding the digital signals up to T2, intermediate addition value A addition value A (intermediate addition value A
T2 Tl Z全  T2 Tl Z all
T2 T1 を全加算値 A で除算)を示す識別値を算出する識別値算出部 26と、 2個のシンチ  T2 T1 is divided by the total added value A).
T2  T2
レータアレイのシンチレータ 19a, 19bで発光した発光パルスに基づぐ識別値算出 部 26で算出された各シンチレータアレイの識別値間の中間値 Kに対して、識別値算 出部 26で算出された識別値が大きい値力 vj、さい値かを判別する判別部 28とを備え ている。したがって、判別部 28での判別結果により、 γ線検出器 1で検出された γ線 力 2個あるうちのいずれのシンチレータアレイ 18a、 18bへの入射に基づいて検出さ れたものであるかを識別することができる構成となっている。なお、 A /A を算出し Identification value calculation unit 26 for the intermediate value K between the identification values of each scintillator array calculated by the scintillator array 26 based on the emission pulses emitted by the scintillators 19a and 19b of the modulator array A value vj having a large value and a determination unit 28 for determining whether the value is a small value are provided. Therefore, based on the discrimination result in the discriminator 28, it is determined which of the two γ-ray forces detected by the γ-ray detector 1 is detected based on the incident light on the scintillator arrays 18a and 18b. It can be identified. Calculate A / A
Tl T2  Tl T2
、その算出結果が中間値 Kよりも大きい場合は減衰時間の短いシンチレータ 19a、逆 に小さい場合は減衰時間の長いシンチレータ 19bであると識別できる。中間値 Kは、 加算部 25での加算過程で 2パターンの波形の最も離れた時間 Fs X m (Fs X m： Fsは AZD変換のサンプリング間隔、 mは加算回数)での値である A を設定した両波高  If the calculation result is larger than the intermediate value K, it can be identified as a scintillator 19a with a short decay time, and conversely if it is small, it is identified as a scintillator 19b with a long decay time. The intermediate value K is the value A at the time Fs X m (Fs X m: Fs is the sampling interval of AZD conversion, m is the number of additions) of the two patterns of waveforms in the addition process in the adder 25. Set both wave heights
T1  T1
値 (電圧値)の中間値とし、判別を行うためのデータとして、予め実験により取得され たものであり、 FPGA7の ROM9の一機能である中間値データテーブル 27として記 憶されている。なお、上述したシンチレータアレイ識別部 24は、シンチレータアレイ識 別手段に相当する。上述した加算部 25は、加算手段に相当する。上述した識別値 算出部 26は、識別値算出手段に相当する。上述した判別部 28は、判別手段に相当 し、判別処理時に中間値データテーブル 27に記憶された中間値を読み出されるもの である。 It is an intermediate value of the value (voltage value), and is obtained as a data for discrimination in advance by experiment, and is recorded as an intermediate value data table 27 that is a function of ROM9 of FPGA7. It is remembered. The scintillator array identifying unit 24 described above corresponds to a scintillator array identifying unit. The adding unit 25 described above corresponds to adding means. The identification value calculation unit 26 described above corresponds to an identification value calculation unit. The determination unit 28 described above corresponds to a determination unit, and reads the intermediate value stored in the intermediate value data table 27 during the determination process.
[0041] また、位置演算処理部 5a, 5bは、シンチレータアレイ識別部 24で識別されたシン チレータアレイ 18a, 18bに応じて、入射タイミング算出部 4a, 4bで算出された入射タ イミングの補正を行うか否かの判別を行い、判別の結果に基づいて、入射タイミング の補正を行う入射タイミング補正部 29を備えている。具体的には、シンチレ一タァレ ィ識別部 24で識別されたシンチレータアレイのシンチレータが減衰時間の長いシン チレータ 19bであると識別された場合、入射タイミング算出部 4a, 4bで算出された入 射タイミング t を t — Δ t (入射タイミング補正値）とする演算処理を行 ヽ、 FPGA7の  In addition, the position calculation processing units 5a and 5b correct the incident timings calculated by the incident timing calculation units 4a and 4b according to the scintillator arrays 18a and 18b identified by the scintillator array identification unit 24. An incident timing correction unit 29 is provided for determining whether or not, and correcting the incident timing based on the determination result. Specifically, when the scintillator of the scintillator array identified by the scintillation array identifying unit 24 is identified as the scintillator 19b having a long decay time, the incident timing calculated by the incident timing calculating units 4a and 4b Execute t processing where t is t — Δ t (incident timing correction value)
SR SR SR SR
RAM10の一機能である補正後入射タイミング記憶部 30に出力する。逆にシンチレ ータアレイ識別部 24で識別されたシンチレータアレイのシンチレータが減衰時間の 短いシンチレータ 19aであると識別された場合、入射タイミング t は何も補正されず、 The result is output to the corrected incident timing storage unit 30, which is a function of the RAM 10. Conversely, when the scintillator of the scintillator array identified by the scintillator array identifying unit 24 is identified as the scintillator 19a having a short decay time, the incident timing t is not corrected,
SF  SCIENCE FICTION
シンチレータアレイ識別部 24に入力された入射タイミング t そのままを補正後入射タ  Incident timing t input to the scintillator array identification unit 24
SF  SCIENCE FICTION
イミング記憶部 30に出力する。また、補正後入射タイミング記憶部 30は、入射タイミ ング t と入射タイミング t との関係において、シンチレータアレイ 18aとシンチレータ Output to the imming storage unit 30. In addition, the corrected incident timing storage unit 30 includes a scintillator array 18a and a scintillator in relation to the incident timing t and the incident timing t.
SF SR SF SR
アレイ 18bとで減衰時間が異なることによる検出時間の差が補正された、入射タイミン グ t と入射タイミング t とが一時的に記憶される構成となっている。なお、上述した入 The structure is such that the incident timing t and the incident timing t, in which the difference in detection time due to the difference in attenuation time from the array 18b is corrected, are temporarily stored. The above-mentioned input
SF SR SF SR
射タイミング補正部 29は、入射タイミング補正手段に相当する。  The irradiation timing correction unit 29 corresponds to incident timing correction means.
[0042] なお、この Δ t (入射タイミング補正値）は、補正を行うためのデータとして、予め実 験により取得されたものであり、シンチレータアレイ 18aとシンチレータアレイ 18bとで 異なる立ち上がり時間の時間差を入射タイミング補正値として、 FPGA7の ROM9の 一機能である補正データテーブル 31に記憶されている。なお、入射タイミング補正部 29は、補正処理時に補正データテーブル 31に記憶された入射タイミング補正値を 読み出されるものである。 [0042] This Δt (incidence timing correction value) is obtained in advance by experiment as data for performing correction, and the time difference between rise times different between the scintillator array 18a and the scintillator array 18b is obtained. As an incident timing correction value, it is stored in a correction data table 31 which is one function of ROM9 of FPGA7. The incident timing correction unit 29 reads out the incident timing correction value stored in the correction data table 31 during the correction process.
[0043] 同時計数処理部 6は、補正後入射タイミング記憶部 30に記憶された入射タイミング t , t を一定時間（例えば 128ns)毎に読み取り、 2つ γ線検出器 1からの入射タイミ[0043] The coincidence processing unit 6 includes the incident timing stored in the corrected incident timing storage unit 30. t and t are read at regular intervals (for example, 128 ns), and two incident timings from γ-ray detector 1 are read.
SF SR SF SR
ング t , t 、この場合 4通りの組み合わせを同時計数し、この 4通りの組み合わせに T, t, in this case, four combinations are counted simultaneously, and these four combinations are counted.
SF SR SF SR
より算出された時間差が所定の時間範囲であるタイミングウィンド Tw (例えば、 6ns) 内であれば、有効な同時計数カウントとし、そうでない場合は無効とする構成である。  If the calculated time difference is within a timing window Tw (for example, 6 ns) within a predetermined time range, a valid coincidence count is set. Otherwise, the count is invalid.
[0044] また、タイミングウィンド Twにつ!/、て、シンチレータアレイが 1層（個）である各 γ線 検出器 1に入射された γ線を同時計数処理した場合において、図 7を用いて説明す る。図 7は、タイミングウィンド Twを説明するためのグラフである。縦軸 Aにイベント（ 同時計数処理される y線検出）回数、横軸 Tに同時計数処理された γ線の検出の時 間差とすると、図 7に示すようなタイミングスペクトルが得られる、このタイミングスぺタト ルは、横軸 Τが 0 ( γ線の検出の時間差がない）の場合のイベント回数が多ぐ時間 差が長いほど、イベント回数が減っていぐつまり、横軸 Τが 0の時、縦軸 Αのイベント 回数がピークとなるようなガウス分布に近いグラフとなっている。ここで、図 7に示すタ イミングスペクトルにおける縦軸 Aの中間値 AZ2を半値幅とし、この半値幅を 2倍した 時間範囲がタイミングウィンド Twとなる。  [0044] Further, when the γ-rays incident on each γ-ray detector 1 in which the scintillator array has one layer (pieces) are simultaneously counted for the timing window Tw! explain. FIG. 7 is a graph for explaining the timing window Tw. If the vertical axis A represents the number of events (y-line detection for simultaneous counting) and the horizontal axis T represents the time difference for the detection of γ-rays, the timing spectrum shown in Fig. 7 is obtained. The timing spectrum shows that when the horizontal axis 0 is 0 (there is no time difference in the detection of γ rays), the number of events increases. The longer the time difference, the smaller the number of events. In other words, the horizontal axis の is 0. The graph is close to a Gaussian distribution with the peak number of events on the vertical axis 縦 軸. Here, an intermediate value AZ2 of the vertical axis A in the timing spectrum shown in FIG. 7 is a half width, and a time range obtained by doubling the half width is a timing window Tw.
[0045] ここで、具体的に、同時計数処理部 6での同時計数処理について図 8,図 9を用い て説明する。図 8は、シンチレータアレイの減衰時間が異なることによる検出時間の 差を補正していない場合のタイミングスペクトルを示すグラフである。図 9は、シンチレ ータアレイの減衰時間が異なることによる検出時間の差を補正した場合のタイミング スペクトルを示すグラフである。  Here, the coincidence counting process in the coincidence processing unit 6 will be specifically described with reference to FIGS. FIG. 8 is a graph showing the timing spectrum when the difference in detection time due to the different decay times of the scintillator array is not corrected. Figure 9 is a graph showing the timing spectrum when the difference in detection time due to different decay times of the scintillator array is corrected.
[0046] まず、従来の同時計数処理では、図 8に示すように、シンチレータアレイ 18aとシン チレータアレイ 18bとで減衰時間が異なることによる検出時間の差を補正しないこと から、同時計数処理部 6で行われる 4通りの同時計数は、例えば、 2つの γ線検出器 1の一方の γ線検出器 1を MD1,他方の γ線検出器 1を MD2とすると、 MDlt と  First, in the conventional coincidence processing, as shown in FIG. 8, since the difference in detection time due to the difference in the decay time between the scintillator array 18a and the scintillator array 18b is not corrected, the coincidence processing unit 6 The four simultaneous counts performed are, for example, MDlt and MDlt when one γ-ray detector 1 of two γ-ray detectors 1 is MD1, and the other γ-ray detector 1 is MD2.
SF science fiction
MD2t , MDlt と MD2t , MDlt と MD2t , MDlt と MD2t との 4通りの組Four combinations of MD2t, MDlt and MD2t, MDlt and MD2t, MDlt and MD2t
SF SF SR SR SF SR SR SF SF SR SR SF SR SR
み合わせがあり、この 4通りの組み合わせのうち、 MDlt と MD2t , MDlt と MD  Of these four combinations, MDlt and MD2t, MDlt and MD
SF SF SR  SF SF SR
2t とでは、シンチレータアレイの減衰時間は異ならないことから、時間差は発生せ Since the decay time of the scintillator array is not different from 2t, there is no time difference.
SR SR
ず同時計数において問題はないが、 MDlt と MD2t , MDlt と MD2t とでは、  There is no problem in coincidence counting, but MDlt and MD2t, MDlt and MD2t
SR SF SR SR  SR SF SR SR
シンチレータアレイ 18aとシンチレータアレイ 18bとで減衰時間が異なることによる時 間差が生じ、有効な同時計数カウントとされない場合が発生し、感度が低下すること になる。 When the decay time differs between scintillator array 18a and scintillator array 18b Differences occur, and there are cases where the coincidence count is not valid, resulting in a decrease in sensitivity.
[0047] これに対して、本実施例の場合では、シンチレータアレイ 18aとシンチレータアレイ 18bとで減衰時間が異なることによる検出時間の差を入射タイミング補正部 29により 補正されるので、図 9に示すように、同時計数処理部 6で行われる 4通りの同時計数 は、タイミングウィンド Tw内に入っており、有効な同時計数カウントの数え落としは生 じず、感度が低下しない。なお、このタイミングウィンド Twは、 FPGA7の ROM9の一 機能であるタイミングウィンド記憶部 32に記憶されている。なお、上述した同時計数 処理部 6は、同時計数手段に相当する。上述したタイミングウィンド記憶部 32は、タイ ミンダウインド記憶手段に相当する。  On the other hand, in the case of the present embodiment, the difference in detection time due to the difference in attenuation time between the scintillator array 18a and the scintillator array 18b is corrected by the incident timing correction unit 29, which is shown in FIG. As described above, the four types of coincidence performed by the coincidence processing unit 6 are within the timing window Tw, and the effective coincidence count is not counted down, and the sensitivity does not decrease. The timing window Tw is stored in the timing window storage unit 32 which is a function of the ROM 9 of the FPGA 7. The above-described coincidence processing unit 6 corresponds to coincidence means. The timing window storage unit 32 described above corresponds to a time window storage unit.
[0048] 次に、被検体 Mから放出された γ線が γ線検出器 1に入射され、同時計数処理部 [0048] Next, the γ rays emitted from the subject M are incident on the γ ray detector 1, and the coincidence processing unit
6で γ線が同時計数処理されるまでの動作について図 1,図 3,図 10を用いて順番 に説明する。図 10は、入射タイミング発生カゝら入射タイミング補正処理されるまでのフ ローチャートである。 The operation up to the simultaneous counting process of γ rays in Fig. 6 will be explained in order with reference to Figs. FIG. 10 is a flow chart from the incident timing generation to the incident timing correction processing.
[0049] まず、図 1に示すように、被検体の関心部位力 互いにほぼ 180° 方向に放出され る 2本の γ線が対向する γ線検出器 1のシンチレ一タブロック 15に入射される。 y線 は、図 3示すように、シンチレ一タブロック 15を構成する発光パルスの減衰時間が短 いシンチレータアレイ 18aのシンチレータ 19aと、発光パルスの減衰時間が長いシン チレータアレイ 18bのシンチレータ 19bとのそれぞれで光が生じ、これらの光は、ライ トガイド 16に導かれて、入射した位置（シンチレータ 19a, 19bの X方向と Y方向）に 基づいて、光電子増倍管 (PMT) 17の 4面の PMT光電変換膜に分配されて到達す る。さらに、光電子増倍管 17では、光を電気信号 (アナログ信号）に変換して増幅回 路 2a, 2bに出力される。増幅回路 2a, 2bでは、アナログ信号を電圧増幅させて、入 射タイミング算出部 4a, 4bおよび AZD変 3a, 3bに出力させる。さらに、 AZD 変 3&, 3bに入力されたアナログ信号は、デジタル信号に AZD変換され、 AZ D変換信号記憶部 23に一時的に記憶される。 [0049] First, as shown in FIG. 1, two γ-rays emitted from the subject's region of interest in a direction of approximately 180 ° are incident on the scintillator block 15 of the γ-ray detector 1 facing each other. . As shown in FIG. 3, the y-line represents the scintillator array 18a of the scintillator array 18a with a short decay time of the light emission pulse constituting the scintillator block 15, and the scintillator 19b of the scintillator array 18b with a long decay time of the light emission pulse. Light is generated at the light guide 16 and guided to the light guide 16, and the PMT on the four sides of the photomultiplier tube (PMT) 17 is based on the incident position (X direction and Y direction of the scintillators 19a and 19b). It is distributed to the photoelectric conversion film. Further, in the photomultiplier tube 17, the light is converted into an electric signal (analog signal) and output to the amplification circuits 2a and 2b. In the amplification circuits 2a and 2b, the analog signal is voltage amplified and output to the incident timing calculation units 4a and 4b and the AZD variables 3a and 3b. Further, the analog signal input to the AZD transformations 3 & , 3b is AZD converted into a digital signal and temporarily stored in the AZD conversion signal storage unit 23.
[0050] ここで、被検体 Mから放出された γ線が γ線検出器 1に入射され、同時計数処理 部 6で γ線が同時計数処理されるまでの動作のうち、増幅回路 2a, 2bからのアナ口 グ信号を入力した入射タイミング算出部 4a, 4bで入射タイミング t , t が算出され、 [0050] Here, among the operations until the γ-rays emitted from the subject M are incident on the γ-ray detector 1 and the γ-rays are simultaneously counted by the coincidence processing unit 6, the amplification circuits 2a and 2b Ana mouth from The incident timings t and t are calculated by the incident timing calculators 4a and 4b,
SF SR  SF SR
この入射タイミング t , t が入射タイミング補正部 29で処理されるまでの流れについ  The flow until the incident timing t, t is processed by the incident timing correction unit 29 is as follows.
SF SR  SF SR
て、図 10を用いて説明する。図 10は、光電子増倍管 17から出力された電気信号が 増幅回路 2a, 2bから出力された後の信号処理を示すフローチャートである。  This will be described with reference to FIG. FIG. 10 is a flowchart showing signal processing after the electrical signal output from the photomultiplier tube 17 is output from the amplifier circuits 2a and 2b.
[0051] (ステップ S1)入射タイミング算出部 4a, 4bの ARC20は、増幅回路 2a, 2bからの アナログ信号を入力に基づいて、入射タイミング t , t を算出し、さらに、タイミング発 (Step S1) The ARC 20 of the incident timing calculation units 4a and 4b calculates the incident timings t and t based on the analog signals from the amplifier circuits 2a and 2b, and further generates the timing.
SF SR  SF SR
生回路 21で、これら入射タイミング t , t をデジタル信号に変換し、入射タイミング記  The live circuit 21 converts these incident timings t and t into digital signals and records the incident timings.
SF SR  SF SR
憶部 22に出力する。ここで、入射タイミング記憶部 22に入射タイミング t , t が記憶  Output to memory unit 22. Here, the incident timings t and t are stored in the incident timing storage unit 22.
SF SR  SF SR
(入射タイミング t , t が発生）された場合は、ステップ S 2に進み、入射タイミング t ,  (Incident timing t 1, t is generated), the process proceeds to step S 2 and the incident timing t 1,
SF SR SF  SF SR SF
t が発生されるまでは、次の動作には進まない (ステップ SIの動作を繰り返す)。 Until t is generated, the next operation is not performed (the operation of step SI is repeated).
SR SR
[0052] (ステップ S2)位置演算処理部 5a, 5bのシンチレータアレイ識別部 24は、入射タイ ミング t , t が発生されたことに基づいて、 AZD変換信号記憶部 23に一時的に記 (Step S2) The scintillator array identification unit 24 of the position calculation processing units 5a and 5b temporarily stores the incident timings t and t in the AZD conversion signal storage unit 23 based on the occurrence of the incident timings t and t.
SF SR SF SR
憶されて 、る AZD変換後のデジタル信号を読み出し、ステップ S3に進む。  Remember, read the digital signal after AZD conversion and go to step S3.
[0053] (ステップ S3)位置演算処理部 5a, 5bのシンチレータアレイ識別部 24は、 AZD変 換後のデジタル信号を順次加算することで積分する。途中の T (Fs X m:Fsは AZD (Step S3) The scintillator array identifying unit 24 of the position calculation processing units 5a and 5b integrates the digital signals after the AZD conversion by sequentially adding them. T (Fs X m: Fs is AZD
1  1
変換のサンプリング間隔、 mは加算回数)まで途中加算値 A と、発光終了時点 T (F  During the conversion sampling interval, m is the number of additions) and the intermediate addition value A and the flash end point T (F
T1 2 s X n:nは総加算回数）までの全加算値 A とを取得し、ステップ S4に進む。  T1 2 s X n: n is obtained as a total addition value A up to (total number of additions), and the process proceeds to step S4.
T2  T2
[0054] (ステップ S4)さらに、位置演算処理部 5a, 5bのシンチレータアレイ識別部 24は、 途中加算値 A と^ロ算値 A とから、中加算値 A ロ算値 A を示す識別値を  [0054] (Step S4) Further, the scintillator array identification unit 24 of the position calculation processing units 5a and 5b obtains an identification value indicating the intermediate addition value A and the second calculation value A from the intermediate addition value A and the second calculation value A.
Tl T2 Tl T2  Tl T2 Tl T2
算出し、 2個のシンチレータアレイ 18a, 18bの各シンチレータ 19a, 19bで発光した 発光パルスに基づぐ中間値データテーブル 27に記憶された中間値 Kに対して、識 別値が大きい値力 vj、さい値かを判別する。さらに、これら判別された結果を示す信号 を位置演算処理部 5a, 5bの入射タイミング補正部 29に出力する。ここで、識別値が 中間値 Kよりも小さい場合はステップ S5に進み、逆に識別値が中間値 Kよりも大きい 場合は、ステップ S6に進む。  The value vj having a large discriminant value relative to the intermediate value K stored in the intermediate value data table 27 based on the light emission pulses emitted from the scintillators 19a and 19b of the two scintillator arrays 18a and 18b. Determine whether the value is small. Further, a signal indicating the determined result is output to the incident timing correction unit 29 of the position calculation processing units 5a and 5b. If the identification value is smaller than the intermediate value K, the process proceeds to step S5. If the identification value is larger than the intermediate value K, the process proceeds to step S6.
[0055] (ステップ S5)位置演算処理部 5a, 5bの入射タイミング補正部 29は、シンチレータ アレイ識別部 24で識別されたシンチレータアレイが、識別値が中間値 Kよりも小さい 場合、つまり、発光パルスの減衰時間が長いとして、入射タイミング算出部 4a, 4bで 算出された入射タイミング t — A t (入射タイミング補正値)とする補正を行うと判 [0055] (Step S5) The incident timing correction unit 29 of the position calculation processing units 5a and 5b determines that the scintillator array identified by the scintillator array identification unit 24 has an identification value smaller than the intermediate value K, that is, a light emission pulse. In the incident timing calculation unit 4a, 4b It is determined that correction is made to the calculated incident timing t — At (incident timing correction value).
SR SR  SR SR
別する。したがって、入射タイミング補正部 29は、入射タイミング記憶部 22に記憶さ れている入射タイミング t を読み出し、この入射タイミング t に t —A tの補正演算  Separate. Therefore, the incident timing correction unit 29 reads the incident timing t stored in the incident timing storage unit 22, and corrects t −A t for this incident timing t.
SR SR SR  SR SR SR
処理を行い、この t - A tを示す信号を補正後入射タイミング記憶部 30に出力し、補  Processing, and outputs a signal indicating this t-At to the corrected incident timing storage unit 30 for correction.
SR  SR
正後入射タイミング記憶部 30で記憶される。  It is stored in the post-front incidence timing storage unit 30.
[0056] (ステップ S6)また、位置演算処理部 5a, 5bの入射タイミング補正部 29は、シンチ レータアレイ識別部 24で識別されたシンチレータアレイ力 識別値が中間値 よりも 大きい場合、つまり、発光パルスの減衰時間が短いとして、入射タイミング算出部 4a, 4bで算出された入射タイミング t の補正を行わないと判別する。したがって、入射タ (Step S6) Further, the incident timing correction unit 29 of the position calculation processing units 5a and 5b determines that the scintillator array force identification value identified by the scintillator array identification unit 24 is larger than the intermediate value, that is, the light emission pulse. It is determined that the incident timing t calculated by the incident timing calculation units 4a and 4b is not corrected. Therefore, the incident tag
SR  SR
イミング記憶部 22に記憶されている入射タイミング t を読み出し、この入射タイミング  Read the incident timing t stored in the imming storage unit 22
SR  SR
t に補正処理を行わず、そのままの入射タイミング t を示す信号を補正後入射タイミ No correction processing is performed on t, and the signal indicating the incident timing t is corrected as is.
SR SR SR SR
ング記憶部 30に出力し、補正後入射タイミング記憶部 30で記憶される。  Output to the storage unit 30 and stored in the corrected incident timing storage unit 30.
[0057] 次に、被検体 Mから放出された γ線が γ線検出器 1に入射され、同時計数処理部 6で γ線が同時計数処理までの動作のうち、この入射タイミング t , t が入射タイミン [0057] Next, γ-rays emitted from the subject M are incident on the γ-ray detector 1, and the incident timing t 1, t 2 of the operations up to the coincidence counting processing by the coincidence processing unit 6 is Incident time
SF SR  SF SR
グ補正部 29で処理されてから同時計数処理までの流れについて、図 7を用いて説明 する。同時計数処理部 6は、補正後入射タイミング記憶部に記憶された入射タイミン グ t , t を 128ns毎に読み取り、 2つ γ線検出器 1からの入射タイミング t , t 、この The flow from the processing by the correction unit 29 to the coincidence counting process will be described with reference to FIG. The coincidence processing unit 6 reads the incident timings t and t stored in the corrected incident timing storage unit every 128 ns, and the incident timings t and t from the two γ-ray detectors 1 and
SF SR SF SR SF SR SF SR
場合 4通りの組み合わせを同時計数し、この 4通りの組み合わせにより算出された時 間差がタイミングウィンド Tw (例えば、 6ns)内であれば、有効な同時計数カウントとし 、そうでない場合は無効とする。  In case 4 combinations are counted simultaneously, if the time difference calculated by these 4 combinations is within the timing window Tw (e.g. 6ns), it is a valid coincidence count, otherwise it is invalid. .
[0058] 上述した核医学診断装置によれば、入射タイミング補正部 29で、シンチレ一タァレ ィ識別部 24で識別されたシンチレータアレイに応じて、入射タイミング算出部 4a, 4b で算出された入射タイミングの補正を行うか否かの判別を行い、判別の結果に基づ いて、入射タイミングの補正を行うので、発光パルスの減衰時間の異なるシンチレ一 タアレイ 18aとシンチレータアレイ 18bを用いて同時計数した場合にお!、ても、この発 光パノレスの減衰時間の異なることにより生じるシンチレータアレイ 18aとシンチレータ アレイ 18bとにおける検出の時間差を補正により解消することができる。したがって、 Ύ線検出器 1において、発光パルスの減衰時間の異なるシンチレータを用いた場合 でも、高感度であり、再構成画像の劣化を生じさせず、正確な断層画像を得ることが できる。 According to the nuclear medicine diagnostic apparatus described above, the incident timing calculated by the incident timing calculation units 4a and 4b in the incident timing correction unit 29 according to the scintillator array identified by the scintillation tally identification unit 24. In case of simultaneous counting using scintillator array 18a and scintillator array 18b with different emission pulse decay times, the incident timing is corrected based on the determination result. However, the detection time difference between the scintillator array 18a and the scintillator array 18b caused by the difference in the decay time of the light emitting panel can be eliminated by correction. Therefore, when scintillators with different emission pulse decay times are used in the shoreline detector 1. However, it is highly sensitive, and an accurate tomographic image can be obtained without causing degradation of the reconstructed image.
[0059] また、シンチレータアレイ識別部 24の加算部 25で順次加算することに基づいて、 識別値算出部 26で算出された識別値が大きい値力 vj、さい値かを判別することができ る。つまり、シンチレータアレイ識別部 24は、シンチレータで発光した発光ノ《ルスが 何れのシンチレータアレイであるかを識別することができる。また、従来積分器で行わ れていた積分動作を加算手段で順次加算する加算動作に置き換えることができるの で、部品点数の削減とコストダウンを図ることができる。  [0059] Further, based on the sequential addition by the addition unit 25 of the scintillator array identification unit 24, it is possible to determine whether the identification value calculated by the identification value calculation unit 26 is a large value force vj or a small value. . That is, the scintillator array identifying unit 24 can identify which scintillator array is the light emitting light emitted by the scintillator. In addition, since the integration operation that has been conventionally performed by the integrator can be replaced with the addition operation in which the addition means sequentially adds, the number of parts can be reduced and the cost can be reduced.
[0060] この発明は、上記実施形態に限られることはなぐ下記のように変形実施することが できる。  [0060] The present invention is not limited to the embodiment described above, and can be modified as follows.
[0061] (1)上述した実施例では、 PET装置を例に採って説明したが、この発明は、放射性 薬剤が投与された被検体から発生した放射線を同時計数して核医学診断を行う核 医学装置であれば、 PET装置に限定されずに適用することができる。  [0061] (1) In the above-described embodiments, the PET apparatus has been described as an example. However, the present invention provides a nucleus for performing nuclear medicine diagnosis by simultaneously counting radiation generated from a subject to which a radiopharmaceutical is administered. Any medical device can be applied without being limited to the PET device.
[0062] (2)上述した実施例では、 PET装置と X線 CT装置とを備えた PET— CTのように、 核医学診断装置と X線 CT装置とを組み合わせた装置にも適用することができる。  [0062] (2) In the above-described embodiment, the present invention can be applied to an apparatus combining a nuclear medicine diagnostic apparatus and an X-ray CT apparatus, such as a PET-CT equipped with a PET apparatus and an X-ray CT apparatus. it can.
[0063] (3)タイミングウィンド記憶部 32は、複数個のシンチレータアレイそれぞれの組み合 わせに対応するタイミングウィンド Twを記憶するようにしてもよい。したがって、複数 個のシンチレータアレイそれぞれの組み合わせにより異なるタイミングウィンド Twを 用いることで、精度の高い同時計数を行うことができ、偶発同時計数や散乱同時計 数などの影響を減らし、ノイズの少な 、高画質な画像を得ることができる。  [0063] (3) The timing window storage unit 32 may store a timing window Tw corresponding to each combination of the plurality of scintillator arrays. Therefore, by using different timing windows Tw depending on the combination of multiple scintillator arrays, it is possible to perform highly accurate coincidence counting, reduce the effects of accidental coincidence counting, scattered clock counting, etc. A high-quality image can be obtained.
[0064] (4)タイミングウィンド記憶部 32は、複数個のシンチレータそれぞれの組み合わせ に対応するタイミングウィンド Twを記憶するようにしてもよい。したがって、複数個の シンチレータそれぞれの組み合わせにより異なるタイミングウィンド Twを用いることで 、精度の高い同時計数を行うことができ、偶発同時計数や散乱同時計数などの影響 を減らし、ノイズの少な 、高画質な画像を得ることができる。  [0064] (4) The timing window storage unit 32 may store a timing window Tw corresponding to each combination of a plurality of scintillators. Therefore, by using different timing windows Tw depending on the combination of multiple scintillators, it is possible to perform highly accurate coincidence, reduce the effects of accidental coincidence and scatter coincidence, reduce noise and achieve high image quality. An image can be obtained.
[0065] (5)上述した実施例では、位置演算処理部 5a, 5bの入射タイミング補正部 29は、 入射タイミング算出部 4a, 4bで算出された入射タイミング t を t — A t (入射タイミン  (5) In the above-described embodiment, the incident timing correction unit 29 of the position calculation processing units 5a and 5b uses the incident timing t calculated by the incident timing calculation units 4a and 4b as t — At (incident timing).
SR SR  SR SR
グ補正値）とする演算処理を行い、また、入射タイミング t を補正処理せず、入射タイ ミング t —A tと、入射タイミング t とを補正後入射タイミング記憶部 30に記憶するよCorrection processing), and the incident timing t is not corrected and the incident timing t Ming t —A t and incident timing t are stored in the corrected incident timing storage unit 30.
SR SF SR SF
うにしていたが、位置演算処理部 5a, 5bの入射タイミング補正部 29は、入射タイミン グ算出部 4a, 4bで算出された入射タイミング t を t + A tとする演算処理を行い、ま  However, the incident timing correction unit 29 of the position calculation processing units 5a and 5b performs calculation processing with the incident timing t calculated by the incident timing calculation units 4a and 4b as t + At.
SF SF  SF SF
た、入射タイミング t を補正処理せず、入射タイミング t + A tと入射タイミング t とを  The incident timing t is not corrected, and the incident timing t + At and the incident timing t are
SR SF SR  SR SF SR
補正後入射タイミング記憶部 30に記憶するようにしてもょ 、。  It may be stored in the corrected incident timing storage unit 30.
[0066] (6)上述した実施例では、シンチレータアレイ 18a、 18bで異なる立ち上がり時間の 時間差を入射タイミング補正値として、補正データテーブル 31に記憶するようにして いた力 シンチレータ 19間で異なる立ち上がり時間の時間差を入射タイミング補正値 として、補正データテーブル 31に記憶するようにしてもよい。したがって、シンチレ一 タ 19間で異なる立ち上がり時間の時間差を補正することができ、さらに、検出感度を 向上させることができる。  [0066] (6) In the embodiment described above, the time difference between the rise times different in the scintillator arrays 18a and 18b is stored in the correction data table 31 as the incident timing correction value. The time difference may be stored in the correction data table 31 as the incident timing correction value. Therefore, it is possible to correct the time difference between the rise times that differ between the scintillators 19, and to further improve the detection sensitivity.
[0067] (7)上述した実施例では、シンチレータアレイ 18a、 18bで異なる立ち上がり時間の 時間差を予め入射タイミング補正値として、補正データテーブル 31に記憶するように していたが、識別値算出手段で算出された識別値から、例えば識別値を変数とした 簡単な一次関数で入射タイミング補正値をリアルタイムに算出するようにしてもよい。  (7) In the above-described embodiment, the time difference between the rise times different in the scintillator arrays 18a, 18b is stored in advance in the correction data table 31 as the incident timing correction value. From the calculated identification value, for example, the incident timing correction value may be calculated in real time with a simple linear function using the identification value as a variable.
[0068] (8)上述した実施例では、シンチレ一タブロック 15は、 γ線入射側（前段）に発光パ ルスの減衰時間が短いシンチレータ 19aとして Lu Y SiO (LYSO)を用いたシン  (8) In the above-described embodiments, the scintillator block 15 is a scintillator block using Lu Y SiO (LYSO) as the scintillator 19a having a short emission pulse decay time on the γ-ray incident side (front stage).
X 2-X 5  X 2-X 5
チレータアレイ 18aと、ライトガイド 16側（後段）に発光パルスの減衰時間が長いシン チレータ 19bとして Ce濃度 0. 5molの Gd SiO (GSO)を用いたシンチレータアレイ 1  A scintillator array 18a and a scintillator array using 0.5 mol of Ce-concentrated Gd SiO (GSO) as the scintillator 19b on the light guide 16 side (rear stage) with a long decay time of the emission pulse 1b
2 5  twenty five
8bとで構成するようにしていた力 シンチレ一タブロック 15を構成するシンチレータァ レイ 18aのシンチレータ 19aと、シンチレータアレイ 18bのシンチレータ 19bとは、 Ce 濃度 0. 5molの Gd SiO (GSO) , Ce濃度 1. 5molの Gd SiO (GSO) , Lu SiO (L  The scintillator array 18a scintillator 19a and the scintillator array 18b scintillator 19b consist of a Ce concentration of 0.5 mol Gd SiO (GSO) and a Ce concentration. 1. 5mol Gd SiO (GSO), Lu SiO (L
2 5 2 5 2 5 2 5 2 5 2 5
SO) , Lu Gd SiO (LGSO) , Lu Y SiO (LYSO) , Bi Ge 0 (BGO) , Nal, B SO), Lu Gd SiO (LGSO), Lu Y SiO (LYSO), Bi Ge 0 (BGO), Nal, B
X 2-X 5 X 2-X 5 4 3 12 aF , CsFを選択し、種々組合わせて使用しても良い。  X 2-X 5 X 2-X 5 4 3 12 aF and CsF may be selected and used in various combinations.
2  2
[0069] (9)上述した実施例では、シンチレ一タブロック 15は、シンチレータアレイ 18aとシ ンチレータアレイ 18bとの 2層（個）組合わせたものとして説明した力 2層（個）以外の 複数層（個）であってもよい。また、各シンチレータに備えるシンチレータ 19a, 19bの 数を 8 X 8本として説明したが、これ以外の複数本備えるようにしてもよ!/、。 (10)上述した実施例では、受光素子を光電子増倍管 17として説明したが、これ以 外の受光素子、例えば、フォトダイオードやアバランシェフオトダイオードなどを用い てもよい。 [0069] (9) In the embodiment described above, the scintillator block 15 has a force other than two layers (pieces) described as a combination of two layers (pieces) of the scintillator array 18a and the scintillator array 18b. Multiple layers (pieces) may be used. In addition, although the number of scintillators 19a and 19b provided in each scintillator is described as 8 × 8, a plurality of other scintillators may be provided! (10) In the above-described embodiments, the light receiving element is described as the photomultiplier tube 17, but other light receiving elements such as a photodiode or an avalanche photodiode may be used.

Claims

請求の範囲 The scope of the claims
[1] 複数個のシンチレータを 2次元的に密着配置し、 γ線入射深さ方向に発光パルス の減衰時間が異なる複数個のシンチレータアレイを光学的に結合したシンチレータ ブロックと、前記シンチレ一タブロックで発光した発光ノ ルスを電気信号に変換する 受光素子と、前記受光素子力 出力された電気信号について、前記シンチレータァ レイに入射したタイミングを算出する入射タイミング算出手段と、前記受光素子から出 力された電気信号が複数個あるうちのいずれのシンチレータアレイに入射したかを識 別するシンチレータアレイ識別手段と、前記シンチレータアレイ識別手段で識別され たシンチレータアレイに応じて、前記入射タイミング算出手段で算出された入射タイミ ングの補正を行うか否かの判別を行い、判別の結果に基づいて、入射タイミングの補 正を行う入射タイミング補正手段と、を備えて ヽることを特徴とする核医学診断装置。  [1] A scintillator block in which a plurality of scintillators are two-dimensionally arranged closely and a plurality of scintillator arrays in which the emission pulse decay times differ in the γ-ray incident depth direction are optically coupled, and the scintillator block A light-receiving element that converts the light-emitting noise emitted from the light-receiving element into an electric signal; an incident-timing calculating means that calculates a timing at which the electric signal output from the light-receiving element force is incident on the scintillator array; and an output from the light-receiving element The scintillator array identifying means for identifying which of the plurality of electric signals incident to the scintillator array and the scintillator array identified by the scintillator array identifying means are calculated by the incident timing calculating means. To determine whether or not to correct the incident timing, and based on the result of the determination A nuclear medicine diagnostic apparatus, comprising: an incident timing correction unit that corrects the incident timing.
[2] 請求項 1に記載の核医学診断装置において、前記受光素子から出力された電気 信号であるアナログ信号をデジタル信号に変換する AZD変換器を備え、前記シン チレータアレイ識別手段は、前記 AZD変 で変換されたデジタル信号を順次カロ 算する加算手段と、前記加算手段において、前記シンチレ一タブロックで発光した発 光パルスの発光開始時力 発光終了時までの途中である途中時点までのデジタル 信号を加算した途中加算値および前記シンチレ一タブロックで発光した発光パルス の発光開始時力 発光終了時までのデジタル信号を加算した []算値から、途中加 算値を全加算値で除算した値を示す識別値を算出する識別値算出手段と、前記識 別値算出手段で算出された各シンチレータアレイの識別値間の中間値に対して、前 記識別値算出手段で算出された識別値が大きい値力 vj、さい値かを判別する判別手 段と、を備えていることを特徴とする核医学診断装置。  [2] The nuclear medicine diagnosis apparatus according to claim 1, further comprising: an AZD converter that converts an analog signal that is an electric signal output from the light receiving element into a digital signal, and the scintillator array identifying means includes the AZD converter. In addition means for sequentially caloring the digital signal converted in step (b), and in the addition means, a digital signal up to the middle point in the middle of the light emission start power of the light emission pulse emitted by the scintillator block until the end of light emission. The value obtained by adding the halfway addition value and the power at which the light emission pulse emitted from the scintillator block emits light to the digital signal until the end of light emission. The value obtained by dividing the intermediate addition value by the total addition value. The identification value calculation means for calculating the identification value indicating the difference between the identification values of the scintillator arrays calculated by the identification value calculation means and the identification value calculation means. Nuclear medicine diagnosis apparatus characterized in that it comprises identification values calculated by means large value power vj, and determination means to determine again value.
[3] 請求項 1に記載の核医学診断装置にお！ヽて、前記入射タイミング補正手段で補正 した入射タイミングおよび前記入射タイミング補正手段で入射タイミングを補正しない と判別された入射タイミングを用いて同時計数を行う同時計数手段と、前記同時計数 手段での同時計数が同時と判定される所定の範囲を示すタイミングウィンドを前記複 数個のシンチレータアレイそれぞれの組み合わせに対応するタイミングウィンドとして 記憶するタイミングウィンド記憶手段とを備えていることを特徴とする核医学診断装置 [3] The nuclear medicine diagnostic apparatus according to claim 1! The simultaneous counting means for performing simultaneous counting using the incident timing corrected by the incident timing correcting means and the incident timing determined not to correct the incident timing by the incident timing correcting means, and the simultaneous counting means by the simultaneous counting means A nuclear medicine diagnosis, comprising: a timing window storing means for storing a timing window indicating a predetermined range in which counting is determined simultaneously as a timing window corresponding to a combination of each of the plurality of scintillator arrays. apparatus
[4] 請求項 1に記載の核医学診断装置にお!ヽて、前記入射タイミング補正手段で補正 した入射タイミングおよび前記入射タイミング補正手段で入射タイミングを補正しない と判別された入射タイミングを用いて同時計数を行う同時計数手段と、前記同時計数 手段での同時計数が同時と判定される所定の範囲を示すタイミングウィンドを前記複 数個のシンチレータそれぞれの組み合わせに対応するタイミングウィンドとして記憶 するタイミングウィンド記憶手段と、を備えて!/ヽることを特徴とする核医学診断装置。 [4] In the nuclear medicine diagnosis apparatus according to claim 1, using the incident timing corrected by the incident timing correcting means and the incident timing determined not to correct the incident timing by the incident timing correcting means. A timing window for storing simultaneous counting means for performing simultaneous counting and a timing window indicating a predetermined range in which simultaneous counting by the simultaneous counting means is determined to be simultaneous as timing windows corresponding to combinations of the plurality of scintillators; A nuclear medicine diagnostic apparatus characterized by comprising a storage means!
[5] 請求項 1に記載の核医学診断装置において、前記シンチレ一タブロックと前記受光 素子とを光学的に結合するライトガイドを備えていることを特徴とする核医学診断装 置。  5. The nuclear medicine diagnosis apparatus according to claim 1, further comprising a light guide that optically couples the scintillator block and the light receiving element.
[6] 請求項 1に記載の核医学診断装置において、複数個のシンチレータアレイは、 Ce 濃度 0. 5molの Gd SiO (GSO) , Ce濃度 1. 5molの Gd SiO (GSO) , Lu SiO (L  [6] In the nuclear medicine diagnosis apparatus according to claim 1, the plurality of scintillator arrays include Ce concentration of 0.5 mol Gd SiO (GSO), Ce concentration of 1.5 mol Gd SiO (GSO), Lu SiO (L
2 5 2 5 2 5 2 5 2 5 2 5
SO) , Lu Gd SiO (LGSO) , Lu Y SiO (LYSO) , Bi Ge 0 (BGO) , Nal, B SO), Lu Gd SiO (LGSO), Lu Y SiO (LYSO), Bi Ge 0 (BGO), Nal, B
X 2-X 5 X 2-X 5 4 3 12 aF , CsFのいずれかのシンチレータにより構成されていることを特徴とする核医学診 X 2-X 5 X 2-X 5 4 3 12 Nuclear medicine diagnosis characterized by a scintillator of either aF or CsF
2 2
断装置。  Cutting device.
[7] 請求項 1に記載の核医学診断装置において、前記受光素子は光電子増倍管であ ることを特徴とする核医学診断装置。  7. The nuclear medicine diagnostic apparatus according to claim 1, wherein the light receiving element is a photomultiplier tube.
[8] 請求項 1に記載の核医学診断装置において、前記受光素子はフォトダイオードで あることを特徴とする核医学診断装置。 8. The nuclear medicine diagnostic apparatus according to claim 1, wherein the light receiving element is a photodiode.
[9] 請求項 1に記載の核医学診断装置において、前記受光素子はアバランシェフオト ダイオードであることを特徴とする核医学診断装置。 9. The nuclear medicine diagnostic apparatus according to claim 1, wherein the light receiving element is an avalanche photodiode.
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JPWO2007043137A1 (en) 2009-04-16
CN101208616B (en) 2011-07-27

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